Non-invasive method and device to monitor cardiac parameters

ABSTRACT

A method of and a device for non-invasively measuring the hemodynamic state of a subject or a human patient involve steps and units of non-invasively measuring cardiac cycle period, electrical-mechanical interval, mean arterial pressure, and ejection interval and converting the measured electrical-mechanical interval, mean arterial pressure and ejection interval into the cardiac parameters such as Preload, Afterload and Contractility, which are the common cardiac parameters used by an anesthesiologist. 
     The converted hemodynamic state of a patient is displayed on a screen as a three-dimensional vector with each of its three coordinates respectively representing Preload, Afterload and Contractility. Therefore, a medical practitioner looks at the screen and quickly obtains the important and necessary information.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates to a non-invasive method and device tomonitor cardiac parameters.

2. Description of the Prior Art

At the present time anesthetics, (drugs which induce loss of sensation)are often used for surgical operations. A general anesthetic generallycauses a progressive depression of the central nervous system andinduces the patient to lose consciousness. In contrast, a localanesthetic affects sensation at the region where it is applied.

Generally, prior to the operation, the patient is anesthetized by aspecialized medical practitioner (“anesthesiologist”) who administersone or more volatile liquids or gases such as nitrous oxide, halothane,isoflurane, sevoflurane, desflurane, and etc. Alternatively,non-volatile sedative-hypnotic drugs such as pentothal, propofol, andetomidate may are administered by injection or intravenous infusion.Opioid analgesics like morphine, fentanyl, or sufenanil may likewise beadministered by injection or infusion, to relieve pain by raising thepain sensation threshold.

Some of the objectives of a correctly administered general anestheticare as follows: Firstly, the patient should be sufficiently anesthetizedso that his/her movements are blocked. If the patient's movements arenot sufficiently blocked, the patient may begin to “twitch” (involuntarymuscle reflexes) during the operation, which may move or disturb theoperating field that is an area being operated. Such blockage ofmovement occurs with a paralysis of the central nervous system after thesensory cortex is suppressed. The paralysis sequentially affects thebasal ganglia, the cerebellum and then the spinal cord. The medulla,which controls respiratory, cardiac and vasomotor centers, is depressedby the anesthetic in a dose dependent fashion. When respiration iscompletely depressed by the anesthetic, it must be performed for thepatient by the anesthesiologist, using either a rubber bag, or automaticventilator.

Secondly, the patient should be sufficiently unconscious so as to feelno pain and be unaware of the operation. Patients have sued for medicalmalpractice because they felt pain during the operation or were aware ofthe surgical procedure. Once unconsciousness has been achieved, powerfuldepolarizing and non-depolarizing muscle relaxant drugs can be given toassure a quiescent undisturbed operating field for the surgeon.

Thirdly, the anesthesia should not be administered in an amount so as tolower blood pressure to the point where blood flow to the brain may bereduced to a dangerous extent to cause cerebral ischemia and hypoxia.The dangerous extent is generally below 50 mm Hg for mean arterialpressure (MAP). For example, if the blood pressure is too low for over10 minutes, the patient may not regain consciousness. This criticalpressure will vary with the patient's medical condition. In patientswith hypertension, for example, the critical pressure below which injurycan occur, is elevated.

A skilled anesthesiologist may monitor the vital signals such asbreathing, heart rate, and blood pressure of the patient to determine ifmore or less anesthetic is required. Often, the anesthesiologist looksinto the patient's eyes to determine the extent of the dilation of thepupils as an indication of the level or depth of the effect of theanesthesia. The depth is also called “plane of anesthesia.” However,there may be a number of problems with such complete reliance on theskill and attention of the anesthesiologist. In modern practice, theeyes are frequently taped shut to avoid abraision or ulceration of thecornea of the eye. Since some operations may be prolonged for 10 to 15hours, the attention of the anesthesia nurse or anesthesiologist mayflag or fail. Therefore, it is important to provide a simple method tomonitor the patient's state of the cardiovascular system.

The state or performance of the cardiovascular system can be describedin terms of hemodynamic parameters. One such parameter is the cardiacoutput (CO). Much effort has been invested in non-invasive methods tomeasure the CO. (See Klein, G., M.D., Emmerich, M., M.D., ClinicalEvaluation of Non-invasive Monitoring Aortic Blood Flow, (ABF) by aTransesophageal Echo-Doppler-Device. Anesthesiology 1998; V89 No. 3A:A953; Wallace, A W., M.D, Ph.D., et.al., Endotracheal Cardiac OutputMonitor, Anesthesiology 2000; 92:178-89). But the cardiac output is justa summary parameter or a final common result of many possiblehemodynamic states. In clinical practice, fluid administration andvasoactive drug infusion therapy are not directed to changing the CO perse. Rather, they are directed to the CO's component parameters such asthe heart rate (HR) and the Stroke Volume (SV). The relation among theHR, the SV and the CO is given by

CO=HR[SV]  Eq. 1

The SV, in turn, is a function of three constituent parameters. ThePreload (P) measures the “tension” in cardiovascular muscle at enddiastole. The Afterload (A) measures the “resistance” to the bloodoutflow from the left ventricle. The Contractility (C) measures the rateof rising of the “strain” in cardiovascular muscle. SV increases withincreasing P and C and decreases with increasing A. (See Braunwald, E.,M.D., ed., Heart Disease, A Textbook of Cardiovascular Medicine, FourthEdition, Philadelphia, W.B. Saunders Company, 1992, p. 420). In otherwords, the following relation holds.

SV=f(P,A,C)  Eq. 2

where f( )is a predetermined function.

One way of looking at Eq. 2 is to understand that SV is a function of avector in a three dimensional space. This vector is just (P,A,C). Theaxes of the vector space are mutually perpendicular and include P, A,and C. By Eq. 1, CO is linearly proportional to SV by the factor of HR.We can therefore understand that HR is scalar and operates on a vectorin a three dimensional, hemodynamic vector space, H. Substituting Eq. 2in Eq. 1, we have

CO=HR[f(P,A,C)]  Eq. 3

Every possible hemodynamic state in a given system is represented by aunique point in the (P,A,C) space and is scaled by HR. There is a subsetof points within H, that are compatible with life. The subject is aphysiologic hemodynamic vector subspace that we can call P. P is whollycontained in H. If we can track the position of the hemodynamic vectorin this hemodynamic vector space, that is, follow its trajectory, thenwe can have fairly complete knowledge of what the effects ofpharmacologic and fluid therapy are during the perioperative period. Wecan titrate fluids and diuretics, pressors and afterload reducers,anesthetics, inotropes and negative inotropes against a change in theposition of the vector and its relative projection onto each of thethree mutually perpendicular axes.

Preload, Afterload, and Contractility have been traditionally assessedby invasive methods. Preload has been approximated by PulmonaryCapillary Wedge Pressure (PCWP), which is measured with a Swan-Ganzpulmonary artery balloon-tipped catheter that is wedged into thepulmonary arterial circulation. Preload has also been approximated bymeasuring the area of the left ventricle image at end-diastole with 2-Dechocardiography. Afterload has been approximated using the Swan-Ganzcatheter to perform thermodilution cardiac output measurements, andmeasurements of Mean Arterial Pressure (MAP) and Central Venous Pressure(CVP) to calculate the Systemic Vascular Resistance. This is done inanalogy with Ohm's law for electrical resistance. In clinical practice,Contractility is approximated as the cardiac ejection fraction. Thisrequires the methods of nuclear medicine or 2D echocardiography.Alternatively, Contractility is approximated as the maximum rate of riseof left ventricular pressure (P) in systole. This is just the maximumvalue of the first derivative of pressure with respect to time duringsystolic ejection. That is, the approximation is dP/dt max. (SeeBraunwald, E., M.D., ed., Heart Disease, A Textbook of CardiovascularMedicine, Fourth Edition, Philadelphia, W.B. Saunders Company, 1992, p.431). Measuring dP/dt max requires catheterization of the leftventricle. This hazardous and arrythmogenic procedure is usuallyreserved for the cardiac catheterization lab.

Swan-Ganz catheters are invasive. Invasion is the occasion of clinicalmischief. Most experienced clinicians understand this in a visceral way.Pulmonary artery rupture, hemo-pneumothorax, pulmonary infarcts,bacterial endocarditis, large vein thrombosis, and intraventricularknotting are just a few of the well-known complications that result fromusing this device. Some authors have advocated a moratorium on theiruse, believing that the risks outweigh the benefits. (see Connors, A. F.Jr., M.D., et. al., The Effectiveness of Right Heart Catheterization inthe Initial Care of the Critically Ill Patients, J. Amer. Med. Assn.,1996; 276:889-897; Dalen, J. E., Bone R. C.: Is It Time to Pull thePulmonary Catheter? J. Amer. Med. Assn., 1996; 276:916-8). 2-Dtransesophageal echocardiography devices are prohibitively expensive.They also require specialized image interpretation skills. They arestill minimally invasive. Likewise, the methods of Nuclear Medicine areexpensive, requiring a cyclotron to produce specializedradiopharmaceuticals and specialized image interpretation skills.Moreover, Nuclear Ejection Fractions cannot be done continuously and inreal time. They can be used to assess baseline cardiac function. Theycannot be used to titrate fluid therapy and drug infusions from momentto moment.

Newer technologies have emerged such as the Hemosonic device from ArrowInternational (see Klein, G., M.D., Emmerich, M., M.D., ClinicalEvaluation of Non-invasive Monitoring Aortic Blood Flow, (ABF) by aTransesophageal Echo-Doppler-Device. Anesthesiology 1998; V89 No. 3A:A953). This minimally invasive device uses a trans-esophageal Dopplerplaced in the esophagus and one-dimensional A-mode echocardiograph. TheDoppler measures velocity of blood in the descending aorta while theA-mode ultrasound is used to measure the descending aortic diameter inreal time. Integrating blood velocity times aortic diameter over theejection interval gives the stroke volume. Stroke volume times heartrate gives cardiac output. Dividing cardiac output into the MeanArterial Pressure gives Systemic Vascular Resistance. Measuring peakblood acceleration gives Contractility. Because the device measuresblood flow in the descending aorta, it ignores blood flow to the headand both arms. Thus, it ignores about 30% of the total cardiac outputand cannot measure Preload. Because the device sits in the thoracicesophagus, it cannot be used on people who are awake.

If it were possible to approximate Preload, Afterload, and Contractilityusing non-invasive means or equipment which is already ubiquitous andrelatively inexpensive, then many more patients on whom invasivemonitors and 2D echocardiography devices are not currently used, couldbenefit from hemodynamic monitoring without its high costs and highrisks. This possibility includes many pediatric patients, renalpatients, pregnant patients, and cardiac patients presenting fornon-cardiac surgery. The above described non-invasive hemodynamicmonitoring on a beat-to-beat basis would represent a great improvementin the state of the art, resulting in significant reductions in the costof care and in perioperative morbidity.

There is a need for a low-cost, low risk, non-invasive metric againstwhich a wide array of cardiovascular support drug administrations andinfusions can be adjusted, in order to optimize the condition ofpatients with a wide variety of cardiovascular medical conditions,within the constraints of said conditions and illnesses. Because of itslow-cost, low-risk character, it should render possible the non-invasiveclinical monitoring of a wide range of cardiovascular illness, in theoperating room and intensive care unit, and also from locations outsidethe traditional operating room theater and critical care units. Itshould allow clinicians to pinpoint and quantify the component causes ofacute decompensations in chronic cardiovascular illness, and to use thisinformation to modify therapy in such a way as to prevent frequent,costly hospitalization.

Accordingly, there is a need to provide apparatuses and methods forcontinuously and accurately providing real-time information relating tocardiac output in the form of volume blood flow based upon non-invasivemeasurements. There is also a need to provide apparatuses and associatedmethods for monitoring cardiac output which results in a reducedlikelihood of an undetected catastrophic event. Additionally, there is aneed to provide a method for monitoring cardiac output in which the riskof infection is eliminated or significantly reduced.

Accordingly, it is an objective of the present invention to providedevices and methods for detecting, assessing the cardiac timing of,grading, and diagnosing a variety of vascular and arrythmia conditions.

It is another objective of the present invention to provide devices andmethods for non-invasively monitoring the hemodynamic state of a patientand providing approximate information on the Preload, Afterload andContractility.

It is to be noted that the scope of this invention is not simply in thesphere of anesthesia, but in the totality of medicine, includingoutpatient, ambulatory, and critical care medicine. For example, itsolves the problem of optimizing fluid administration and the use ofdiuretics and inotropes (like digitalis) and afterload reducers, likethe vasodilator Captopril, in patients with Congestive Heart Failure(CHF). Too little fluid, and the cardiac output becomes insufficient toperfuse vital organs like the brain, heart, and kidney, resulting inorgan failure and death. Too much fluid, and the pumping capacity of thecompromised left heart is overwhelmed, allowing fluid to back up intothe lungs, causing a diffusion barrier to oxygenation. Fluid welling upin the lungs effectively causes the patient to drown. In thiscircumstance, patients need to be hospitalized, intubated, andventilated in an ICU. By adjusting the diuretic dose against thePreload, or its analogue, and by adjusting the Digitalis dose againstthe contractility, and adjusting the Captopril dose against the SVR orits analogue, you can keep someone with CHF out of the hospital forlonger periods of time, saving both money and grief.

SUMMARY OF THE INVENTION

In a first aspect, the present invention provides a method of monitoringcardiac parameters. The method of monitoring the cardiac parametersincludes the steps of: non-invasively measuring a plurality ofpredetermined non-invasive cardiac parameters from a subject; andconverting the non-invasive cardiac parameters into a plurality ofinvasive cardiac analogues based upon a set of predetermined conversionequations.

In a second aspect, the present invention provides a system formonitoring cardiac parameters. The system for monitoring the cardiacparameters includes: a non-invasive cardiac parameter measuring unit fornon-invasively measuring a plurality of predetermined non-invasivecardiac parameters from a subject; and a conversion unit connected tothe non-invasive cardiac parameter measuring unit for converting thenon-invasive cardiac parameters into a plurality of invasive cardiacanalogues based upon a set of predetermined conversion equations.

In a third aspect, the present invention provides a system forretrofitting existing non-invasive cardiac parameter measuring devicesto generate invasive cardiac analogues. The system for retrofitting theexisting non-invasive cardiac parameter measuring devices includes: aninterface unit for receiving predetermined non-invasive cardiacparameters of a subject from the existing non-invasive cardio monitoringdevices; and a conversion unit connected to the interface unit forconverting the non-invasive cardiac parameters into a plurality of theinvasive cardiac analogues based upon a set of predetermined conversionequations.

In a fourth aspect, the present invention provides a method ofdetermining a patient's cardiac contractility. The method of determiningthe patient's cardiac contractility includes the steps of:non-invasively measuring the patient's electrocardiograph having apredetermined electrical wave; determining a first point having aminimum within a predetermined cardiac cycle based upon thepredetermined electrical wave; non-invasively measuring the patient'sarterial pressure with respect to time; determining a second point inthe predetermined cardiac cycle when a second derivative of apredetermined physiological function with respect to time reachesmaximum; and obtaining the cardiac contractility based upon the firstpoint and the second point.

In a fifth aspect, the present invention provides a system fordetermining a patient's cardiac contractility. The system fordetermining a patient's cardiac contractility includes: anelectrocardiogram unit for non-invasively measuring the patient'selectrocardiograph having a predetermined electrical wave; an arterialpressure measuring unit for non-invasively measuring the patient'sarterial pressure with respect to time; and a determination unitconnected to the electrocardiogram unit and the arterial pressuremeasuring unit for determining a first point having a minimum within apredetermined cardiac cycle based upon the predetermined electrical waveand for determining a second point in the predetermined cardiac cyclewhen a second derivative of a predetermined physiological function withrespect to time reaches maximum based upon the patient's arterialpressure, the determination unit obtaining the cardiac contractilitybased upon the first point and the second point.

In a sixth aspect, the present invention provides a method of monitoringan ischemic event. The method of monitoring the ischemic event includesthe steps of: non-invasively measuring a plurality of predeterminednon-invasive cardiac parameters from a subject; and converting thenon-invasive cardiac parameters into a single invasive cardiac analogueindicative of the ischemic event based upon a predetermined conversionequation.

In a seventh aspect, the present invention provides a system formonitoring an ischemic event. The system for monitoring the ischemicevent includes: a measuring unit for non-invasively measuring aplurality of predetermined non-invasive cardiac parameters from asubject; and a converting unit connected to the measuring unit forconverting the non-invasive cardiac parameters into a single invasivecardiac analogue indicative of the ischemic event based upon apredetermined conversion equation.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a diagram illustrating one preferred embodiment of the devicefor non-invasively monitoring a patient's cardiac parameters accordingto the present invention.

FIG. 2 is a diagram illustrating the first embodiment of the display ofa hemodynamic state according to the present invention.

FIG. 3A is a diagram illustrating the second embodiment of the displayof a hemodynamic state according to the present invention.

FIG. 3B is a diagram illustrating the third embodiment of the display ofa hemodynamic state according to the present invention.

FIG. 4A is a diagram illustrating a first preferred embodiment of thesystem for performing telemedicine according to the current invention.

4B is a diagram illustrating a second preferred embodiment of the systemfor performing telemedicine according to the current invention.

FIG. 5 is a graph illustrating an invasive hemodynamic vector space fora first subject according to the first embodiment of the presentinvention.

FIG. 6 is a graph illustrating a non-invasive hemodynamic vector spacefor the first subject according to the first embodiment of the presentinvention.

FIG. 7 is a graph illustrating an invasive hemodynamic vector space fora second subject according to the first embodiment of the presentinvention.

FIG. 8 is a graph illustrating a non-invasive hemodynamic vector spacefor the second subject according to the first embodiment of the presentinvention.

FIG. 9 is a graph illustrating (Q-A″max) Interval vs. (Q-A) Interval forthe first subject according to the first embodiment of the presentinvention.

FIG. 10 is a graph illustrating average left ventricular ejectioninterval outflow rate as a function of the E-M interval for the firstsubject according to the first embodiment of the present invention.

FIG. 11 is a graph illustrating average left ventricular ejectioninterval outflow rate as a function of the E-M interval for the secondsubject according to the first embodiment of the present invention.

FIG. 12 is a graph illustrating stroke volume as a function of ejectioninterval and E-M interval for the first subject according to the firstembodiment of the present invention.

FIG. 13 is a graph illustrating stroke volume as a function of ejectioninterval and E-M interval for the second subject according to the firstembodiment of the present invention.

FIG. 14 is a graph illustrating average systolic outflow rate vs.Contractility for the first subject according to the first embodiment ofthe present invention.

FIG. 15 is a graph illustrating average systolic outflow rate vs.Contractility for the second subject according to the first embodimentof the present invention.

FIG. 16 is a graph illustrating 1/(Q-A″max) the first subject accordingto the first embodiment of the present invention.

FIG. 17 is a graph illustrating 1/(Q-A″max) the second subject accordingto the first embodiment of the present invention.

FIG. 18 is a graph illustrating [(Ejection Interval)*(Mean ArterialPressure)*(Q-A″max)] as a function of left ventricular end-diastolicpressure for the first subject according to the first embodiment of thepresent invention.

FIG. 19 is a graph illustrating [(Ejection Interval)*(Mean ArterialPressure)*(Q-A″max)] as a function of left ventricular end-diastolicpressure for the second subject (Experiments 5-15) according to thefirst embodiment of the present invention.

FIG. 20 is a graph illustrating [(Ejection Interval)*(Mean ArterialPressure)*(Q-A″max)] as a function of left ventricular end-diastolicpressure for the second subject (Experiments 1-4) according to the firstembodiment of the present invention.

FIG. 21 is a graph illustrating [(Ejection Interval)*(Mean ArterialPressure)*(Q-A″max)] as a function of left ventricular end-diastolicpressure for the second subject (All Experiments) according to the firstembodiment of the present invention.

FIG. 22 is a graph illustrating [(Mean Arterial Pressure)*(Q-A″max)] asa function of Systemic Vascular Resistance for the first subjectaccording to the first embodiment of the present invention.

FIG. 23 is a graph illustrating [(Mean Arterial Pressure)*(Q-A″max)] asa function of Systemic Vascular Resistance for the second subjectaccording to the first embodiment of the present invention.

FIG. 24 is a graph illustrating the Sigmoidal relation between thefilling interval (T−EI) and LVEDP for the first subject according to thesecond embodiment of the present invention.

FIG. 25 is a graph illustrating the Sigmoidal relation between thefilling interval (T-EI) and LVEDP for the second subject according tothe second embodiment of the present invention.

FIG. 26 is a graph illustrating LVEDP in terms of Period, EI, MAPc, andQ−A″max for the first subject according to the second embodiment of thepresent invention.

FIG. 27 is a graph illustrating LVEDP in terms of Period, EI, MAPc, andQ−A″max for the second subject according to the second embodiment of thepresent invention.

FIG. 28 is a graph illustrating the linear relationship between(T-EI)*MAP*(Q−A″max) and exp(LVEDP) for the first subject according tothe second embodiment of the present invention.

FIG. 29 is a graph illustrating the linear relationship between(T-EI)*MAP*(Q−A″max) and exp(LVEDP) for the second subject according tothe second embodiment of the present invention.

FIG. 30 is a graph illustrating SVRc in terms of MAPc and Q-A″max forthe first subject according to the second embodiment of the presentinvention.

FIG. 31 is a graph illustrating SVRc in terms of MAPc and Q-A″max forthe second subject according to the second embodiment of the presentinvention.

FIG. 32 is a graph illustrating an invasive hemodynamic vector space forthe first subject according to the second embodiment of the presentinvention.

FIG. 33 is a graph illustrating a non-invasive hemodynamic vector spacefor the first subject according to the second embodiment of the presentinvention.

FIG. 34 is a graph illustrating a first embodiment of an invasivehemodynamic vector space for the second subject according to the secondembodiment of the present invention.

FIG. 35 is a graph illustrating a first embodiment of a non-invasivehemodynamic vector space for the second subject according to the secondembodiment of the present invention.

FIG. 36 is a graph illustrating a second embodiment of the invasivehemodynamic vector space for the second subject according to the secondembodiment of the present invention.

FIG. 37 is a graph illustrating a second embodiment of the non-invasivehemodynamic vector space for the second subject according to a secondembodiment of the present invention.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

Left Ventricular End-Diastolic Pressure (LVEDP), Systemic VascularResistance (SVR) and the Maximum Rate of Rise of Left VentricularPressure (dp/dtmax) are respectively clinically useful indices andinvasive cardiac analogues of or approximations to Preload, Afterloadand Contractility. Even though these respective pairs of cardiacparameters are not perfectly linear with respect to one another, theyare monotonically increasing with respect to each other. Therefore,LVEDP, SVR and dP/dtmax are also cardiac parameters that are responsiveto cardiac medicines such as fluids and diuretics, pressors andafterload reducers, anesthetics, inotropes and negative inotropes. Thatis precisely why clinicians can rely on LVEDP, SVR and dp/dtmax toadminister the proper dosage of medicines for further controlling theseparameters and therefore for adjusting the state of hemodynamics of thepatient.

In addition, it is an accepted tenet of physiology that a completedescription of the functional state of the heart is given by fourparameters. They are the heart rate, the LVEDP, the SVR, and dP/dtmax.(See Braunwald, E., M.D., ed., Heart Disease, A Textbook ofCardiovascular Medicine, Fourth Edition, Philadelphia, W.B. SaundersCompany, 1992, p. 374-82). The last three of these, which determine thestroke volume, has been typically obtained only at the cost of invasionof the patient.

By making the appropriate substitutions for Preload, Afterload, andContractility, we can rewrite Eqs. 2 and 3 respectively as,

SV=f(LVEDP,SVR,dP/dtmax)  Eq. 4, and

CO=HR[f(LVEDP,SVR,dP/dtmax)]  Eq. 5

Therefore, the state of a hemodynamic system is substantially describedbased upon the above four parameters. Three of these parametersconstitute a vector in a three-dimensional vector space, H′. The axes ofH′ are LVEDP, SVR, and dP/dt max with appropriate units. A function ‘f’of this vector determines the stroke volume, SV. The fourth parameter,the heart rate HR operates linearly as a scalar on the vector todetermine the cardiac output, CO.

In a first aspect, the present invention provides a non-invasive methodof monitoring a patient's first plurality of cardiac parametersresponsive to medicines, measuring non-invasively a second plurality ofcardiac parameters and converting the second plurality of cardiacparameters into the first plurality of cardiac parameters that aredirectly responsive to external medicines.

In a preferred embodiment of the present invention, the first pluralityof cardiac parameters are LVEDP, SVR, and dP/dtmax, which are directlyresponsive to cardiac medicines such as fluids and diuretics, pressorsand afterload reducers, anesthetics, inotropes and negative inotropes.More preferably, the first plurality of cardiac parameters furtherinclude heartrate (HR). The second plurality of cardiac parameters isnon-invasively measured directly using proper instrumentation. Thesecond plurality of non-invasively measured parameters includes meanarterial pressure (MAP), the Ejection Interval (EI) andElectrical-Mechanical Interval (E-M). More preferably, the secondplurality of non-invasively measured parameters further includes HeartRate (HR), which together with MAP, EI and E-M substantially gives acomplete description of the function state of the heart.

EI is the time interval during which systolic ejection takes place. Itstarts when the aortic valve opens and ends when it closes. If anordinary Doppler ultrasound device is placed over the suprasternal notchnear the ascending aorta, inspection of the frequency vs. time curvewill yield the EI. Also, since the time from mitral valve closure toaortic valve opening in systole is small compared to the EjectionInterval, the interval from the first heart sound to the second heartsound measured using a stethoscope or phonocardiogram is optionally auseful approximation to the EI.

E-M is defined by the time between two specific events, an electricalevent and a mechanical event. The electrical event is an eventdetectable on the EKG, which initiates ventricular contraction. Theelectrical event can be the Q-wave, the R-wave or the S-wave. In eachcase, Q, R or S is respectively defined as a point in time when theQ-wave, the R-wave or the S-wave reaches a particular point such as amaximum, a minimum or other predetermined point on the wave. The eventis optionally a ventricular pacing spike.

In some arrhythmias, like ventricular tachycardia with a pulse, there ISno Q-wave (or R-wave, or S-wave). Therefore, another embodiment of ‘E’of the E-M interval is to look at the EKG waveform defining ventriculardepolarization, differentiate it twice with respect to time, and definethe point in time at which the electrical depolarization waveaccelerates maximally upward as ‘E’. This would allow for the definitionof an E-M interval in those instances where there is no recognizable Q,R, or S wave, i.e. when the patient is in extremis. For instance, inventricular tachycardia, the waveform looks like a rapid sine wave. Thisalternative embodiment of ‘E’ may also turn out to be a practically moreaccurate way to determine E-M by more accurately defining ‘E’, to withinnarrower tolerances. The main point, is to find and accurately define aphysiologically identical time point in all possible EKG ventriculardepolarization cycles, which are then compared to one another to createconsistent usable E-M intervals.

The mechanical event is a palpable consequence of ventricularcontraction. It is related to the electrical event and lags theelectrical event in time. The upstroke of the arterial trace from anindwelling arterial catheter qualifies as a mechanical event, and sodoes the instant at which the upward acceleration of the arterialpressure trace is at maximum. In other words, a mechanical event occursat the instant of maximum value of the second derivative of pressurewith respect to time. If the arterial blood pressure (ABP) is given byA(t), then the mechanical event is given by A″(t)max or A″max forsimplicity. Therefore, in one embodiment, the E-M interval (E-M) isfurther defined as Q-A″max, R-A″max or S-A″max.

If we place a Doppler device over a major artery such as the ascendingthoracic aorta near the sternal notch, then the instant of flow velocityupstroke with the onset of systole qualifies as a mechanical event. IfDoppler detected flow is given by F(t), then the instant at which theacceleration in flow is maximum or F″(t) max is also, a usefulmechanical event. Therefore, in another embodiment, the E-M interval(E-M) is defined as Q-F″(t)max, R-F″(t)max or S-F″(t)max.

A useful mechanical event is also obtained from the upstroke of theoptical plethysmographic curve using a pulse oximeter placed on apatient's finger, toe, nose or earlobe. Similarly, the instant ofmaximum upward acceleration of the plethysmographic curve (PM(t)) is aclinically useful mechanical event. In one embodiment, the mechanicalevent is defined as the instant at which the PM(t) curve hits a minimumprior to the detection of flow. Alternatively, the mechanical event isdefined as the instant at which PM(t) curve accelerates maximally upwardas flows become rapid. Differentiating the PM(t) curve twice withrespect to time give us the PM″(t). The instant at which

PM″(t) reaches a maximum value, following the Q-wave (or itssubstitutes) defines a Q−PM″(t)max interval, which is a furtherembodiment of the E-M interval (E-M).

The onset of the first heart sound, representing the closure of themitral valve optionally likewise serves as a useful mechanical event.The instant of maximum amplitude of the first heart sound is optionallyused as a mechanical event as well. It matters little which event isused to define the E-M interval according to the current invention. Byanalogy, the E-M interval is like the interval between a flash oflightning and a clap of thunder. It matters only that we use the sameone consistently when making comparative judgments.

A particular mechanical event is detected using a physiologic sensordeveloped at Empirical Technologies Corp to define the E-M interval.This technology uses a fiberoptic device that sits over the radialartery and vibrates with the arrival of the pulse wave. The vibration ofthe fiberoptic element due to the arterial pulse wave affects thetransmission of a beam of light inside.

Another embodiment detects the mechanical event by placing afiberoptical seismometer device over a large artery to measure thedisplacement of the arterial wall transverse to the direction of bloodflow. The displacement of the arterial wall transverse to the directionof blood flow with respect to time t is defined as TD(t). By analogy, anE-M interval is defined as Q-TD″(t)max. TD″(t)max is the time whenTD″(t), which is the double differentiate of TD(t), reaches its maximumvalue.

Using the interval between the trough of the Q-wave on EKG and theupstroke of the arterial pressure wave in a major artery, the Q-Ainterval (one type of E-M interval), the quantification of myocardialContractility was first described in a letter to the editor of theLancet by Jackson, in 1974. (see Jackson, D. M., M.D, A SimpleNon-Invasive Technique for Measuring Cardiac Contractility, [Letter].Lancet 1974; ii:1457). Using human volunteers, he plotted the decreasein the Q-A interval from baseline at one-minute intervals, whileinfusing isoproteranol. As the infusion came to equilibrium, hedescribed a linear decrease in the Q-A interval with respect to time. Hethen doubled the rate of the infusion and obtained a further lineardecrease in the Q-A interval over time. Of interest, at the lower rateof isoproteranol infusion, the Q-A interval significantly decreased incomparison to baseline while the heart rate changed relatively little.This showed that the decreased Q-A interval was due to an increase inthe inotropic state of the myocardium and not due to an increase in theheartrate. He also described a positive correlation between dP/dtmax andthe decrease in the Q-A interval in anesthetized beagle dogs with leftventricular catheters. He affirmed this correlation using five differentagents, all of which have an effect on the inotropic state of themyocardium, thiopental, calcium, isoproteranol, norepineprine anddigitalis.

In another letter to the editor of the Lancet two months later, Rodbard(see Rodbard, S., Measuring Cardiac Contractility, [Letter]. Lancet1975; I: 406-7) indicated that he had used Jackson's approach for atleast a decade earlier particularly in the diagnosis and evaluation ofhyperthyroid and hypothyroid states. Rodbard described the measurementof the interval from the Q-wave to the Korotkoff sound over a majorartery, the Q-Korotkoff interval (Q-K interval) as well as using aDoppler ultrasound device placed over a major artery to generate aDoppler frequency shift versus time curve (D(t)) to measure the Q-D(t)interval (or Q-D interval).

In contrast, according to the present invention, a more preferredmechanical event is defined by D″(t)max, the time t at which D″(t)reaches maximum value following the peak of the Q-wave or itssubstitute. Similarly, D″(t) is derived by differentiating D(t) twiceagainst time t.

In general, by differentiating a physiologic function M such as A(t),PM(t), F(t), TD(t) or D(t) twice to obtain the time of a usefulmechanical event, an improved accuracy of E-M is achieved. In apreferred embodiment, the mechanical event of E-M is defined as E-M″max,where M″max is defined at the time when M″, which is obtained by doubledifferentiating the physiologic function M against time t, reaches aparticular maximum.

The shorter the E-M interval is, the greater the Contractility of themyocardium becomes. The relation between Q-A, Q-K or Q-D interval andContractility or dp/dtmax has long been in the public domain. (seeCambridge, D., Whiting, M., Evaluation of the Q-A interval as an Indexof Cardiac Contractility in Anesthetized Dogs: Responses to Changes inCardiac Loading and Heart Rate. Cardiovascular Research 1986; 20:444-450). However, as will be disclosed later, the E-M interval is notonly correlated with Contractility but also is used to correlate withother cardiac parameters which are responsive to medicines.

In summary, cardiac output and cardiac state of a patient are correlatedto HR, EI, MAP and E-M, which are the second plurality of cardiacparameters that are non-invasively measured in a direct manner.Therefore, we arrive the following equations, where CO is linear to HR.

CO=HR[f(EI,MAP,E-M)],  Eq. 6

SV=f(EI,MAP,E-M)  Eq. 7

The above relations are mathematically and logically equivalent to therelations among the invasively measured quantities (P, A, C) or itsequivalents (LVEDP, SVR, dP/dtmax).

A first three-dimensional non-invasive vector space M with threemutually perpendicular axes EI, MAP and E-M is constructed even thoughit is not directly responsive to the external medicines. For every pointin the invasive hemodynamic vector space H′, there exists exactly onecorresponding point in the non-invasive hemodynamic vector space M.Moreover, every point in the non-invasive hemodynamic vector space M hasan image in the invasive hemodynamic vector space H′. In the language oflinear algebra, there is a mathematical mapping from the non-invasivehemodynamic vector space M to the invasive hemodynamic vector space H′in a ‘one-to-one’ and corresponding manner. Therefore, in one aspect,the present invention demonstrates there is a one-to-one correlationbetween the non-invasive hemodynamic vector space M and the invasivehemodynamic vector space H′.

A particular hemodynamic state vector in the (EI, MAP, E-M) space doesnot directly show the equivalents or analogues of the invasiveparameters, such as (P, A, C) or (LVEDP, SVR, dP/dtmax). In order to getto an analogue vector in the (P, A, C) or (LVEDP, SVR, dP/dtmax) spacefrom the non-invasively measured vector in the (EI, MAP, E-M) space, apredetermined transformation on the (EI, MAP, E-M) vector is needed.Therefore, the first aspect of the present invention is directed to acorrelation between the above described two vectors or a method ofconverting a vector in the (EI, MAP, E-M) space into an equivalentvector in the (P, A, C) or (LVEDP, SVR, dP/dtmax) space. This conversionmay be implemented in many different forms such as a computer programresiding on a computer.

In one preferred embodiment of the transformation method according tothe present invention, the transformation is accomplished by multiplyingthe (EI, MAP, E-M) vector by a diagonal matrix as shown below. Let x bea vector in the non-invasive hemodynamic space M of the form (EI, MAP,E-M). Let A be the diagonal matrix shown below. If we represent xvertically as a column vector, we can multiply it by the matrix A suchthat Ax=b, where b is a vector of the form ((EI*MAP*E-M), (MAP*E-M),1/(E-M)), that is approximately equivalent to (LVEDP, SVR, dP/dtmax),and a first embodiment of the first plurality of cardiac parametersresponsive to external medicines as being demonstrated in the equationbelow.

${\begin{pmatrix}{M\; A\; P*\left( {E - M} \right)} & 0 & 0 \\0 & {E - M} & 0 \\0 & 0 & {1/\left( {E - M} \right)^{2}}\end{pmatrix}\begin{pmatrix}{E\; I} \\{M\; A\; P} \\\left( {E - M} \right)\end{pmatrix}} = \begin{pmatrix}{{E\; I*M\; A\; {P\left( {E - M} \right)}},} \\{{M\; A\; P*\left( {E - M} \right)},} \\{1/\left( {E - M} \right)}\end{pmatrix}$

The above operation of multiplying the vector by a matrix linearlytransforms the vector x into the vector b. Vector b constitutes a newvector space N or a second Non-invasive Space whose axes are responsiveto external medicines as being verified below. The three mutuallyperpendicular axes of the vector space N are EI*MAP*E-M, MAP*E-M, and1/E-M. The first axis, (EI*MAP*E-M) is linearly proportional to theLVEDP to a first approximation. The second axis, (MAP*E-M) is linearlyproportional to SVR to a first approximation. The third axis, (1/E-M) islinearly proportional to the natural logarithm of dP/dtmax orln(dP/dtmax) to a first approximation. These relations are summarized asfollows:

LVEDP=k1(EI*MAP*E-M)+c1  Eq. 8

SVR=k2(MAP*E-M)+c2  Eq. 9

ln(dP/dt)max=k3(1/E-M_+c3  Eq. 10

Solving Eq. 10 for dP/dt max,

dP/dtmax=Z[exp(k3/E-M)], where Z=exp(c3)  Eq. 11

where k1, k2, k3, and c1, c2, c3 are empirical proportionalityconstants.

Eqs. 8 through 11 are true only to a first approximation. That isbecause while the left hand members of Eqs. 8 through 10 do increasemonotonically with respect to the right hand members, the increases maynot be perfectly linear with respect to one-another. As the patientdeviates further from the physiologic norm, the size of thenon-linearity increases. This is because the relations between the leftand right hand members of Eqs. 8 through 10 are more subtly exponentialthan linear. So within an arbitrarily large neighborhood of a givenphysiologic point, the tangent to the subtle exponential curve gives areasonably good approximation. However, since they are monotonicallyincreasing with respect to one-another, they are practically useful incontrolling proper medicine administration. (EI*MAP*E-M), (MAP*E-M) and(1/E-M) are used to judge the changes in the Preload, Afterload, andContractility due to fluid and drug administration.

In addition to generating a data stream of hemodynamic state vectorsdescribing Preload, Afterload, and Contractility on a beat-to-beatbasis, the method of the present invention also yields a similar datastream about Stroke Volume, SV. SV is a function of only two of thenon-invasive quantities, the Ejection Interval, (EI) and the E-Minterval (E-M). In other words, the following equation expresses therelation

SV=f(EI,E-M)  Eq. 12

Let the average rate of outflow of blood from the Left Ventricle duringthe ejection interval be Fei in cc/sec. Then by definition, thefollowing relation exists.

Fei=SV/EI  Eq. 13

Based on the experimental results disclosed in the present invention,Fei is empirically and linearly proportional to the transcendentalnumber e^(1/E-M). The quantity 1/E-M is the time rate at whichelectromechanical transduction and elastic propagation of the pulse waveor analogous mechanical events occur. So we can write,

Fei=k4*exp(1/E-M)+c4  Eq. 14

Where k4 and c4 are empirical proportionality constants.Solving Eq. 13 for SV, we have

SV=EI*Fei  Eq. 15

Substituting for Fei using Eq. 14, Eq. 15 becomes

SV=EI*[k4*exp(1/E-M)+c4]  Eq. 16 or

SV∝EI*[exp(1/E-M)]  Eq. 16a

Where “∝” means “proportional to.” There are alternative formulations ofSV such as the length or norm of the vector sum of two orthogonalvectors. One of the two orthogonal vectors is a function of EI, and theother is a function of (E-M).

In a second embodiment, the diastolic filling interval (DI) is used toreplace EI. The correlation is improved between (DI, MAP, E-M), which isof the second plurality of non-invasively measured cardiac parametersand (LVEDP, SVR, dP/dtmax) or (P, A, C), which is the first plurality ofinvasive cardiac parameters in a second embodiment. In diastole, theleft ventricular pressure is an exponential function of left ventricularvolume, and this relation holds at any point during the diastolicfilling interval including end-diastole. Therefore, LVEDP is anexponential function of Left Ventricular End Diastolic Volume (LVEDV).

To a reasonable approximation,

DI=T−EI  Eq. 17

where T is the time period of the cardiac cycle. T is easily obtained ina non-invasive manner by measuring the time interval between R-waves inthe EKG and is linearly proportional to the reciprocal of the heartrate, HR in beats per minute. That is,

T=(1/HR)*60 sec/min  Eq. 18

The above approximation ignores the time required for isovolumiccontraction and relaxation. However, since the two intervals arerelatively small fractions of any cardiac cycle, the approximation isuseful.

A more accurate measure of DI is optionally obtained using a 1 MHzDoppler ultrasound device placed on the surface of the patient's chestjust over the left ventricle.) Diastolic filling has a characteristiclow velocity blood flow that causes an analogously low Doppler frequencyshift. The duration of the characteristic low frequency Doppler shiftsubstantially serves as an accurate measure of DI. DI starts when themitral valve opens, and it ends when the mitral valve slams shut. Anordinary stethoscope or phonocardiogram generally indicates when DI endsas marked by the first heart sound, the ‘lub’ of the two soundslub-dub'. In patients with certain pathology, an opening snap of themitral valve is audible in the stethoscope. Perhaps a phonocardiogramshows when the mitral valve opens in most patients. Alternatively, theabove described fiberoptic sensor that is placed upon the precordium ofthe chest serves as a low cost ‘seismometer’ to measure the duration ofthe low frequency vibrations by diastolic filing in the amplitude of thefiberoptic light signal. The Doppler device is more expensive but hasthe advantage for obese patients. Therefore, the correlation between(DI, MAP, E-M) and (LVEDP, SVR, dP/dtmax) or (P, A, C) is defined by thefollowing equations in the preferred embodiment:

LVEDP=k1′((T−EI)*MAP*E-M)+c1′  Eq. 19

SVR=k2′(MAP*E-M)+c2′  Eq. 20

ln(dP/dt)max=k3′(1/E-M)+c3′  Eq. 21

where k1′, k2′, k3′, c1′, c2′ and c3′ are constant for a particularpatient.

Other hemodynamic parameters being equal, the longer the time intervalover which the left ventricle fills, the higher its end-diastolic volumeand pressure becomes. That is, the longer DI is, the higher the LVEDPbecomes. If the EI by itself varies in a useful way with LVEDP, this isdue to a law of physiology relating EI to DI in the steady state. EI byitself has no primary causal relation to LVEDP since it is defined bytwo events that occur in the cardiac cycle after the left ventricle hasfinished filling. The quantity DI=T−EI is logically, temporally andphysiologically prior to the LVEDP. EI by itself is logically,temporally and physiologically posterior to LVEDP.

Using the above described correlation, we now have real-time andnon-invasive measures to be used to express Preload, Afterload,Contractility, Stroke Volume, Heartrate, Cardiac Output, and AverageEjection Outflow Rate. From the foregoing equations, it is relativelysimple to derive useful expressions for Left Ventricular EjectionFraction, whose units are dimensionless, Left Ventricular Stroke Work inunits of Joules, and Left Ventricular Power in Watts.

The existence of the correlation between the first plurality of cardiacparameters and the second plurality of cardiac parameters is verified byusing the method of converting the second plurality of non-invasivecardiac parameters into the first plurality of cardiac parameters thatare measured independently and invasively. The methods of measuring thefirst plurality of non-invasive cardiac parameters are well known toperson skilled in the art. The following are exemplary methods.

Using the averaged waveforms, LVEDP is obtained by inspecting of theLVP(t) waveform and looking for the value of LVEDP just prior to therapid increase in LVP due to systole. Contractility is obtained bydifferentiating the LVP(t) curve with respect to time, and recording themaximum value of the first derivative during systolic ejection,dP/dtmax. Afterload, which is approximated by Systemic VascularResistance (SVR) is obtained by the known formula (see Kaplan, J. A.,M.D., Cardiac Anesthesia, Philadelphia, W.B. Saunders Company, 1993, p.63)

$\begin{matrix}{{S\; C\; R} = \frac{\left( {{M\; A\; P} - {C\; V\; P}} \right)*80}{C\; O}} & {{Eq}.\mspace{14mu} 22}\end{matrix}$

where MAP is the Mean Arterial Pressure in mmHg, and CVP is the CentralVenous Pressure in mmHg. CO is the cardiac output in liters/minute. Itis obtained using the thermodilution technique, with a Swan-Ganzcatheter thermistor connected to a digital temperature vs. time curveintegrator. The constant having a value of 80 is used to convertmmHg/(liter/min) into dyne*sec*cm⁻⁵. CVP was recorded by hand from themonitor at each steady state as with MAP. HR, the heart rate/min, isobtained by measuring the period of the averaged EKG, taking itsreciprocal and then multiplying by 60 sec/min. The HR is divided into COto get the Stroke Volume (SV).

To create the non-invasive hemodynamic state vectors, the followingapproach is used. To measure the Ejection Interval, the LVP and ArterialBlood Pressures (ABP) are graphed simultaneously. The ABP trace is movedbackwards in time until the LVP equals the Diastolic Arterial BloodPressure. This moment in time marks the opening of the aortic valve. Thecurves are followed until they intersect again. This latter point marksthe closure of the aortic valve. The Ejection Interval (EI) is just thetime from opening to closing of the aortic valve. The Mean ArterialPressure (MAP) is simply read from the monitor display. Alternatively,the ABP waveform is integrated over the cardiac period, and then theintegral is divided by the period to get MAP, denoted MAPc. It does notmake a significant difference which approach was used. As indicatedpreviously, EI is easily obtained with an acoustic Doppler device placedin the suprasternal notch, over the ascending aorta. MAP is easilyobtained using a blood pressure cuff and a (DINAMAP). These devices areubiquitous and relatively inexpensive.

Therefore, in another aspect, the present invention provides a system tomonitor a patient's cardiac parameters which are responsive to medicineswith the use of a cuff, at least two electrodes from an EKG instrument,means to measure EI, and a processing means to convert thenon-invasively measured cardiac parameters into invasive cardiacparameter analogues which are responsive to medicines.

FIG. 1 is a diagram illustrating one preferred embodiment of the systemfor non-invasively monitoring a patient's cardiac parameters accordingto the present invention. The system 30 in FIG. 1 includes a cuff 32,two electrodes 34 and 36, a Doppler sensor 38 to measure EI and aprocessing unit 40 to process the signals from the cuff 32, theelectrodes 34 and 36, and the Doppler sensor 38. The cuff 32, theelectrodes 34 and 36, and the Doppler sensor 38 are respectivelyconnected to the processing unit 40 via electrical connections 42, 44,46 and 48. The connections 42, 44, 46 and 48 might be the commonelectrical wires. Alternatively, the connections 42, 44, 46 and 48 mightbe wireless connections. The wireless connections such as infraredconnection and microwave connection are well known to person skilled inthe art. The system 30 is optionally manufactured to be portable. Whenthe system 30 is used to measure or monitor a patient, the cuff 32 isattached to the patient's arm or other appropriate body parts while theelectrodes 34 and 36 are attached to the outer skin in the patient'schest area with a predetermined distance between the electrode 34 andthe electrode 36. Furthermore, the Doppler sensor 38 is placed in thesuprasternal notch over the ascending aorta or over the carotid artery.The processing unit 40 controls the frequency of data acquisition andanalysis and outputs the cardiac parameters of the patient. The outputparameters are used by a medical practitioner to determine the patient'scardiac state and performance. That is, the medical practitionerdetermines whether or not any additional medicine is needed, and he/shealso determines the type and amount of necessary medicines.

More preferably, the system of the present invention further includes amonitoring device which displays the output cardiac parameters on ascreen as a three dimensional vector. One preferred embodiment of thedisplay according to the current invention is illustrated by a diagramas shown in FIG. 2. The monitor 52 has a screen 54, which shows a threedimensional space 56 defined by the three dimensional axes 58, 60 and62. The three dimensions 58, 60 and 62 respectively represent Preload,Afterload and Contractility or their equivalents based upon eitherinvasive or non-invasive measurements. All the cardiac parameters areshown in the three dimensional space 56 as a vector 64. The projections66, 68 and 70 of the vector 64 on the axes 58, 60 and 62 respectivelyrepresent the patient's Preload, Afterload and Contractility. The threedimensional graph on the screen allows a clinician to process a greatdeal of hemodynamic information at one glance. The display substantiallyimproves vigilance in cardiovascular monitoring in the perioperativeperiod.

Even more preferably, as shown in FIG. 3A, in another preferredembodiment of the display according to the current invention, the systemdisplays a vector 80 on the screen 82 that represents a ‘safe’ or‘normal’ hemodynamic state or space. For instance, after the patient issedated but before the surgery begins, the safe hemodynamic state isdetermined. By seeing how the vector 94 moves in real time relative tothe norm vector 80, the operator or the clinician easily and visuallyperceives subtle changes in the patient's hemodynamic profile. Thevector 94 is represented in computer graphics as a ray emanating fromthe origin 86. The projection of the vector 94 onto the Preload axis 88,Afterload axis 90, and the Contractility axis 92 are optionally madedistinct in different colors. Likewise, the three components of the normvector 80 are also optionally marked to create a basis of visualcomparison. A parallel vector 96 in a contrasting color is overlaid uponthe hemodynamic state vector 94. The length of the parallel vector 96represents the size of the cardiac output which is the product of thestroke volume and the heart rate.

Optionally, in FIG. 3A, a box or a safety zone 83 is drawn on the screen82 with the center of the box at the end point 81 of vector 80. Eachedge of the box 83 is either parallel or perpendicular to the axes 88,90 and 92. The length of the edges that are parallel to the Preload axis88 represents the safe range of the patient's Preload. By the sametoken, the length of the edges that are parallel to the Afterload axis90 and the Contractility axis 92 respectively represents the safe rangeof the patient's Afterload and Contractility. Therefore, as long as theend point 95 of the vector 94 is within the safe zone box 83, the vitalcardiac parameters are considered to be within a predeterminedacceptable range. On the other hand, if the end point 95 exits the box83, an appropriate action such as infusion of a suitable medicine isneeded so as to cause the end point 95 enter the box 83.

The deviation of the hemodynamic state vector from a physiological normis indicative of an amount of physiological stress. The degree ofphysiological stress or deviation is defined by a vector cross productbetween the ‘Normal’ vector 100 and the Hemodynamic State Vector 102,and a vector 104 represents the vector cross product as shown in FIG.3B. The vector cross product is a product of the length of the ‘Normal’vector 100, the length of the Hemodynamic State Vector 102 and the sineof the angle between the two vectors. It has a direction of a lineperpendicular to the plane that is defined by the original ‘normal’vector 100 and the hemodynamic state vector 102. It also has an up ordown sign relative to the above plane as given by the right hand rule.

The longer the length of the vector cross product 104, the more seriousthe patient's problem is. The length of the vector cross product is justthe square root of the dot product of the vector cross product withitself. This quantity is just a scalar. As this length of the vectorcross product exceeds a set of predetermined thresholds, a correspondingalarm or alert is optionally given to an operator or a clinician.Additional axes involving the oxygen saturation, the end-tidal carbondioxide and the patient's temperature are alternatively combined in realtime to create a multidimensional vector cross product. Other axes couldbe added as needed or as new modalities of monitoring are developed suchas the processed EEG monitor (BIS) to gauge the depth of anesthesia.Obviously, vector spaces in excess of three dimensions are not easilydisplayed on a screen. But the length of the multidimensional vectorcross product is easily displayed and is properly called a continuousVital Function Scale. Arbitrarily large deviations from the norm alertthe clinician to immediately correct the situation before the patient'slife is threatened. The above described displays 2, 3A and 3B afford theclinician more time by providing the relevant cardiac data to rectifythe problem. As the length of the vector cross product increases, theclinician is at least visually alerted as to the level of deviation fromthe norm.

Furthermore a computer program is implemented to quickly point theclinician's attention to which system or component of themultidimensional vector is a source of the problem so as to saveprecious seconds and to allow more time for a critical intervention. Theabove feature has strong implications for patient safety. The abovedisplays 2, 3A and 3B also substantially reduce the level of skillneeded to recognize the problem. In addition to the clinicians, sometechnicians who have not had the benefit of a medical school educationwould quickly be able to understand the significance of the informationin the displays 2, 3A and 3B. The reduced skill level requirement isprecisely because the system does not require arcane anatomic image orphysiological waveform interpretation skills. The above describedpreferred embodiments are likely to be used with a short learning curveby anyone who can read a graph. Such an easy to learn system such as thedisplays 2, 3A and 3B also has implications for lowering the cost ofhealth care.

Depending on how the hemodynamic state vector moves, various vasoactiveagents are brought to bear upon the problem so as to move the patient'shemodynamics back toward the norm. Agents such as phenylephrine,nitroglycerine, nitroprusside, dopamine, dobutamine and esmolol arelikely to help to stabilize sick patients undergoing the highly variablestresses of surgery. Vasoactive drug infusions are currentlyunderutilized because not enough patients have the full metal jacketinvasive cardiovascular monitoring that is needed to benefit from them.The system according to the present invention increases the wider usageof the easily adjustable vasoactive drugs that are now only routinelyused during cardiac surgery.

Furthermore, the real-time hemodynamic data stream is used to controlthe vasoactive infusion pumps and the level of anesthesia itself via anappropriate computer program. For exemplary applications, the monitoringsystem of the present invention is placed at a patient's home and ismade to communicate via the Internet to a website from which his or herphysician downloads the patient's hemodynamic profile. Furthermore, thesystem is optionally made small enough to be worn by the patient. Themonitored information is stored in the wearable system for over a 24hour or longer period. The system of the present invention isalternatively used in the management of out patients with high bloodpressure or congestive heart failure. The system allows the frequentadjustments of cardiac medications without the need for a patient visit.By allowing frequent and rapid dosage adjustments, the system preventspatients with congestive heart failure from being hospitalized for anacute decompensation. The above prevention capability would avoidenormous expenditures associated with hospitalization includingintubation and ventilation of the patients in critical care units andmanagement of their fluids with Swan-Ganz catheters. Alternatively,patients with cardiomyopathy and severe congestive heart failure alsobenefit from the system according to the current invention to adjusthome dobutamine infusions while they await cardiac transplant.

Renal dialysis patients are similarly helped by the above describedtechnology. For example, Preload of a renal dialysis patient at home isremotely monitored by a clinician while in his office via the Internet.The renal dialysis patient's hemodynamic profile is assessednon-invasively during dialysis. The information is used to manage p.o.fluid intake, restriction, and I.V. fluid administration. Theinformation is also used by the renal dialysis patient himself to adjusthis fluid intake in much the same way that a diabetic patient monitorshis glucose level and adjusts his insulin dose and carbohydrate intake.

The measurement system is calibrated against itself, based upon using avector arbitrarily designated as ‘Norm’ by the clinician when heobserves the patient to be in no acute distress. Alternatively, thesystem undergoes a two or three point calibration against invasivemeasurements made by a cardiologist during cardiac catheterization.Catheterization is undergone only by the sickest patients to see if theyare candidates for coronary artery stent placement, angioplasty, valvesurgery, coronary artery bypass grafting or heart transplant. Once thecalibration is made for a particular patient between the invasivemeasures and the non-invasive measures of cardiac function, it isreasonable that the calibration will persist for many years until thepatient's anatomy significantly changes. The use of the currentinvention would radically decrease the cost of ling term management ofpatients with severe cardiac disease who have had cardiac surgery. Theywould not need subsequent invasive measurements to get the hemodynamicinformation which is vital to their management. They can visit theirdoctors via the Internet using the methods such as telemedicine. It isan advantage that this measurement system significantly expands thescope of the patient services via telemedicine.

FIGS. 4A and 4B are diagrams illustrating preferred embodiments of thesystems 120 and 130 for performing telemedicine according to the currentinvention. In FIG. 4A, one preferred embodiment of the system 30according to the current invention as shown in FIG. 1 is connected to amonitoring and controlling unit 112 via the Internet network 110. Forexample, the system 30 is remotely located at a patient's home in orderto non-invasively measure the patient's cardiac parameters while themonitoring and controlling unit 112 is located at a medicalpractitioner's office. Based upon the measurement results, thepractitioner monitors the patient's cardiac conditions and optionallycontrols the system 30 to acquire additional information. In FIG. 4B,the system 30 of the present invention as shown in FIG. 1 is connectedto a monitoring and controlling unit 112 via a remote connection device114 via a dedicated or shared cable, a wireless telecommunication lineor other suitable connection. The system 130 of FIG. 4B operates in asimilar way as the system 120 of FIG. 4A.

The system of the present invention is used to detect an early warningsign of ischemia in a patient. With ischemia, there is a decrease inContractility and a consequent increase in Preload as the heart tries tomaintain cardiac output by moving up on the Starling curve, which a wellknow relationship between the Cardiac Output and Preload. Then, there isa reflex increase in the Afterload. The above attempt allows the body tomaintain perfusion to vital organs in the face of a decreased cardiacoutput. These events form a recognizable profile in time with asignature time course. The system of the present invention likelyrecognizes the above constellation of events in real time by correlatingthe ischemia event with the non-invasively measured parameters andalerts clinicians of the presence and the extent of ischemic events.Among many possible correlations between ischemia and the non-invasivelymeasured parameters, two useful aspects of the correlations areillustrated in the discussions below. Because decreases in Contractilitywith ischemia are prior to EKG changes in time, clinicians have morelead-time to diagnose and treat life threatening ischemic episodespreceding myocardial infarction. Based upon the above early diagnosis,it is useful to have a nitroglycerine drip in hand in suchcircumstances. It is even more useful to have a physiologic real-timemeasurement to provide an end-point to which the nitroglycerine drip istitrated. Ultimately by preventing myocardial infarctions, the currentinvention helps saving the lives of potentially ischemic patients.

The present invention provides a device such as an interface whichretrofits the existing hospital monitor units to monitor cardiacparameters in a non-invasive manner. The retrofit device of the presentinvention includes a means to measure EI and a processing means toconvert the non-invasively measured results into cardiac parameterswhich are responsive to medicines. In one preferred embodiment of theretrofit system according to the current invention, blood pressure cuffand the EKG are used. Some other currently available devices carry outthe non-invasive measurement of blood pressure by measuring theelectromechanically transduced pulse wave at the wrist. Such a devicecould be used in the present system as well to measure MAP. Such adevice is manufactured by MEDWAVE, 4382 Round Lake Road West, St. PaulMinn. 55112-3923, It sits on the wrist, on the radial artery, andnon-invasively creates a pressure vs. time curve, which would yield MAPas well as ‘M’ of the E-M interval. The device is completelynon-invasive, and its output is visually identical to what can beobtained form an indwelling, invasive arterial catheter pressure trace.The non-invasive measurement data is sent to the processing device viathe interface unit for converting the data into a predetermined format.

In a more preferred embodiment, the fiberoptic sensor mentioned above,is placed over a large artery and yields information on MAP based uponusing the amplitude of low frequency components of the signal.Therefore, the fiberoptic sensor is optionally used not only to providea MAP datastream but also to detect the mechanical event M that definesthe E-M interval in time. If the sensor were placed over a large arterynear the heart such as the carotid artery or ascending aorta, durationof the vibrating signal is used to measure the ejection interval, EI.Now the fiberoptic device gives us information on MAP and the event M inthe (E-M) interval. In that case, the only component required to createthe non-invasive hemodynamic state vectors to provide the cardiacparameters which are responsive to medicines is an EKG. In anotheraspect, the present invention provides a cardiac parameters monitoringsystem including a fiberoptic sensor and an EKG. This monitoring packageis so lightweight that it is worn, so inexpensive that it is disposable,and so mechanically sturdy that it is used in the trenches of theemergency room or the battlefield.

In another aspect, the present invention provides a method or system forgraphically summarizing an entire anesthetic by plotting the trajectoryof the non-invasive hemodynamic state vector at regular time intervals.Alternatively, the trajectory is continuously for record-keepingpurposes and review. Using a three-dimensional graph whose axes arenon-invasive analogues of P, A, and C in the x, y, and z axes, theentire course of an anesthetic or an intensive care unit episode, can bereviewed and understood at one glance. Preferably, the method ofgraphically showing and overlaying heartrate information along thenon-invasive hemodynamic trajectory is used. For example, each datapoint further includes heart rate representation in color along acontinuous spectrum from red to blue, wherein blue denotes lowheartrates, while red denotes high heart rates.

In another aspect, the present invention provides a method or system forcreating a non-invasively derived record of the anesthetic management,an intensive care unit stay, or outpatient cardiovascular diseasemanagement interval. In the event of an untoward event or bad outcome, areviewer is empowered to go back and identify the time evolution ofspecific efficient component causes of said untoward event or badoutcome, since the reviewer is now armed with specific time dependentmeasures of (HR, P, A, C) for improving patient safety and quality ofcare by allowing for more precise diagnosis than the present state ofthe art.

In addition to providing useful parametric information on myocardialfunction based only on non-invasive measurements from the patients whorequire acute medical care, all of the above systems are also applicableto the cardiovascular fitness industry, the life insurance industry, andthe health maintenance organization (HMO) industry to monitor theirpatients' or clients' wellness. For example, using the non-invasivemethod described here, it is possible to determine the percentageincrease in cardiac output that occurs, relative to baseline when atrest, following a prescribed, standardized amount of physical work, suchas dictated by the Bruce Protocol, for a patient on a treadmill. Suchinformation, obtained easily and at virtually zero risk to the patient,may have useful predictive value as an index of cardiac reserve, inpredicting mortality following major surgery, or in predicting mortalitydue to cardiovascular illness. In particular, a higher percentageincrease in cardiac output resulting from a given amount of exercise isa higher index of aerobic cardiac fitness. It could be used by theinsurance industry to adjust their life insurance premiums on anindividual basis, without requiring all clients to undergo cardiaccatheterization. Likewise, it could be used by fitness enthusiasts toquantify their progress in the gym, and by the less enthusiastic, toexercise only exactly as much and as often, as is absolutely necessary.To further confirm and verify the robustness of the correlationsillustrated above, experiments were conducted on two subjects. Theprocedure and results of those experiments are discussed below.

Experimentation

Experiments were performed on two female pigs (Pig 1 and Pig 2) onseparate occasions. The pigs were given general anesthesia withIsoflurane, Oxygen and Fentanyl. They were intubated and mechanicallyventilated. The pigs were monitored using an EKG, a femoral arterialline, and a Swan-Ganz Thermodilution Pulmonary Arterial Catheter. Acapnograph was used to measure end-tidal carbon dioxide. An esophagealthermistor was used to measure body temperature. A pulse oximeter wasused to measure oxygen saturation. In addition, a catheter was placedover a needle, through the anterior chest wall, into the Left Ventricle.Fluoroscopic guidance was used to place the catheter. Intravenouscontrast and pressure readings were used to confirm the presence of thecatheter tip in the Left Ventricle.

Three vasoactive agents were used to create a broad range of possiblehemodynamic states. They are dobutamine, to increase Contractility,nitroglycerine to reduce Preload and Afterload, and phenylephrine toincrease Afterload. One agent was infused at a time by a calibratedinfusion pump at varying rates. At each infusion rate, the system wasallowed to come to equilibrium, and a hemodynamic steady state wascreated. In the first pig experiment, 18 different hemodynamic stateswere created with Pig 1. In the second pig experiment, 15 hemodynamicsteady states were created with Pig 2. Once steady state was achieved,the ventilator was turned off to eliminate respiratory variation. Ateach steady state, the data was digitized and recorded on a laptopfloppy disk. For Pig 1, data was averaged over 60 seconds. For pig 2,data was averaged over 30 seconds. At each hemodynamic steady state andover each data acquisition interval, the waveforms were averaged tocreate an “average” waveform that represented the data over theacquisition interval. Each waveform was arbitrarily synchronized tobegin with the S-wave on the EKG. At the conclusion of the dataacquisition interval, 5 separate thermodilution cardiac outputs weremeasured using 10 cc of saline at room temperature. These were averagedto get an average cardiac output, in liters/minute. The protocol fordrug infusion rates is shown in the following table.

Drug Infusion Protocol Experiment # Drug Concentration Rate, cc/hr Pig#1 1 No drug infusion 2 Dobutamine 1000 micrograms/cc 5 3 Dobutamine1000 micrograms/cc 10 4 Dobutamine 1000 micrograms/cc 15 5 Dobutamine1000 micrograms/cc 10 6 Dobutamine 1000 micrograms/cc 20 7 Dobutamine1000 micrograms/cc 30 8 No drug infusion 9 Nitroglycerine 200micrograms/cc 3 10 Nitroglycerine 200 micrograms/cc 10 11 Nitroglycerine200 micrograms/cc 100 12 Nitroglycerine 200 micrograms/cc 200 13Nitroglycerine 200 micrograms/cc 500 14 No drug infusion 15Phenylephrine 40 micrograms/cc 20 16 Phenylephrine 40 micrograms/cc 4017 Phenylephrine 40 micrograms/cc 80 18 Phenylephrine 40 micrograms/cc120 Pig #2 1 No drug infusion 2 No drug infusion. Pig given 1 liter ofcrystalloid for hydration. 3 Phenylephrine 40 micrograms/cc 20 4Phenylephrine 40 micrograms/cc 60 5 Phenylephrine 40 micrograms/cc 90 6Phenylephrine 40 micrograms/cc 200 7 Phenylephrine 40 micrograms/cc 3008 No drug infusion 9 Dobutamine 1000 micrograms/cc 5 10 Dobutamine 1000micrograms/cc 10 11 Dobutamine 1000 micrograms/cc 15 12 Dobutamine 1000micrograms/cc 20 13 Dobutamine 1000 micrograms/cc 30 14 No drug infusion15 Nitroglycerine 200 micrograms/cc 50 16

To create the invasive hemodynamic state vectors, the following approachwas used. Using the averaged waveforms, LVEDP was obtained by inspectionof the LVP(t) waveform to look for the value of LVEDP just prior to therapid increase in LVP due to systole. Contractility was obtained bydifferentiating the LVP(t) curve with respect to time and recording themaximum value of the first derivative during systolic ejection,dP/dtmax. Afterload that is approximated by Systemic Vascular Resistance(SVR) was obtained by the usual formula (see Kaplan, J. A., M.D.,Cardiac Anesthesia, Philadelphia, W.B. Saunders Company, 1993, p. 63)

$\begin{matrix}{{S\; V\; R} = \frac{\left( {{M\; A\; P} - {C\; V\; P}} \right)*80}{C\; O}} & {{Eq}.\mspace{14mu} 22}\end{matrix}$

where MAP is the Mean Arterial Pressure in mmHg, and CVP is the CentralVenous Pressure in mmHg. CO is the cardiac output in liters/minute. Theabove constant having a value of 80 is used to convert mmHg/(liter/min)into dynesec*cm⁻⁵. CVP was recorded by hand from the monitor at eachsteady state as was MAP. The heart rate (HR) per minute was obtained bymeasuring the period of the averaged EKG, taking its reciprocal, andthen multiplying by 60 sec/min. The HR was divided into the CO to getthe Stroke Volume (SV). Based upon the above parameters, the invasivehemodynamic state vectors were derived.

To create the non-invasive hemodynamic state vectors, the followingapproach was used. To measure the Ejection Interval, the LVP andArterial Blood Pressures (ABP, or A(t)) were graphed simultaneously. TheABP trace was moved backwards in time until the LVP equaled theDiastolic ABP. This moment in time marked the opening of the aorticvalve. The curves were followed until they intersected again. The latterpoint marked the closure of the aortic valve. The Ejection Interval, EIis just the time from opening to closure of the aortic valve. MAP wassimply read from the monitor display. Alternatively, the ABP waveformwas integrated over the cardiac period, and then the integral wasdivided by the period to get MAP. It did not make a significantdifference which approach was used. As indicated previously, EI iseasily obtained with an acoustic Doppler device placed in thesuprasternal notch over the ascending aorta. MAP is also easily obtainedby a blood pressure cuff optionally with a DINAMAP, which is a wellknown blood pressure measuring device. The above devices are ubiquitousand relatively inexpensive.

The E-M interval (more specifically Q-A interval) was measured in twoways. In the first pig experiment, the interval between the Q-wave onEKG and the upstroke of the ABP was measured. In the second pig, the Q-Ainterval measurement was not possible due to noise in the diastolicbaseline arterial pressure trace. To compensate for the inability todefine ‘A’ of the Q-A interval, the curve of ABP was smoothed anddifferentiated twice. In the differentiated curve, the moment of maximumupward acceleration in ABP, A″max, clearly rose above the noise. Whenthe same method was used on the data from the first pig, the lineartransformation worked as well or better. While it can be argued that theuse of an indwelling arterial catheter is not ‘non-invasive’, severalinexpensive and non-invasive methods to define the ‘M’ event at the endof the E-M interval exist.

Incidentally, data processing and graphing were done using MicrocalORIGIN. (see Origin, Data Analysis and Technical Graphics, MicrocalSoftware Inc., One Roundhouse Plaza, Northampton, Mass., 01060).

In a first embodiment of the present invention, LVEDP, SVR and dP/dtmaxare correlated to the non-invasively measured results according to theabove Eqs. 8, 9 and 10. The processing results or conclusions based onthe above correlations are described below. There are close correlationsbetween the invasively measured parameters and the non-invasivelyderived parameters based upon the above described correlation methodsaccording to the present invention.

FIG. 5 shows the results measured by the above described invasive methodfor Pig 1 while FIG. 6 shows the results derived from the resultsmeasured by the above described non-invasive method for Pig 1. FIG. 5shows the Invasive Hemodynamic Vector Space, H′ (LVEDP, SVR, dP/dtmax)in pig 1. FIG. 6 shows the same 18 hemodynamic steady states in theNon-Invasive Hemodynamic Vector space, N. The three mutuallyperpendicular axes of N are given as [EI*MAP*(Q-A″max), MAP*(Q-A″max),1/Q-A″max)] and, mapped to [x, y, z]. With Pig 1, there is real homologybetween the trajectory of the vector from the origin to the point inspace that denotes each hemodynamic state in both alternative sets ofaxes. From experiments 1 through 7, Contractility increases while LVEDPtends to decrease with increasing infusions of dobutamine. A ‘crash’occurs between experiments 7 and 8 when the dobutamine is turned offafter very fast infusions. Both the LVEDP and the SVR decrease betweenexperiments 9 through 13 as the nitroglycerine infusion is increased.The SVR and LVEDP steadily increase as the phenylephrine infusion isincreased among the baseline experiment 14 to experiment 18.Interestingly, Contractility varies only a little between experiments 9through 18. No dobutamine was infused between experiments 9 and 18. Allof these events are mirrored from the Invasive Vector Space H′ to theNon-invasive Vector Space N. Empirically, one sees a ‘one-to-one’mathematical mapping between the two vector spaces as described in Eqs.4 through 11.

FIG. 7 shows the results measured by the above described invasive methodabove for Pig 2, while FIG. 8 shows the results derived from the resultsmeasured by the above described non-invasive method for Pig 2. FIG. 7shows the Invasive Vector Space, H′ in Pig 2. Using the vasoactive druginfusions, Pig 2 is taken on a trajectory through Invasive HemodynamicSpace that stops at fifteen distinct steady states. FIG. 8 shows thesame fifteen hemodynamic steady states in the Non-Invasive HemodynamicVector space, N. The three mutually perpendicular axes of N are given as[EI*MAP*(Q-A″max), MAP*(Q-A″max), 1/Q-A″max)], mapped to [x, y, z].LVEDP increases, and SVR decreases with fluid repletion of one liter ofcrystalloid between states 1 and 2. Contractility does not budge withfluid administration. The change is mirrored in the Non-invasive VectorSpace N. As the phenylephrine infusion is gradually increased betweenstates 4 and 7, the SVR and LVEDP dramatically increase. Contractilitymodestly increases. From state 7 to 8, the phenylephrine is turned off,and there is a dramatic decrease in LVEDP and SVR. This is also wellreflected in N. Now the dobutamine is turned on and gradually increasedfrom states 9 to 13. Contractility significantly increases while LVEDPdecreases. This is consistent with the well known functions ofdobutamine as a cardiac medicine. (see Gilman, A., Goodman L., ThePharmacological Basis of Therapeutics, Seventh Ed., New York, Macmillan,1985, p. 163). Between states 13 and 14, there is a crash as thedobutamine is turned off Contractility rapidly decreases as thedobutamine is metabolized while LVEDP goes up.

During all these hemodynamic states of Pig 2, there is the strikinghomology between FIGS. 7 and 8. It is as though the hemodynamic statevectors were moving in two parallel universes. One universe is withinformation obtained with great risk, cost, and considerable pain whilethe other universe is with information obtained with no risk, low cost,and painlessly.

To further demonstrate the validity of the correlations that are derivedby the preferred method of the present invention, the correlations areverified in the following figures and discussions. FIG. 9 represents the(Q-wave) to (femoral arterial pressure wave, upward accelerationmaximum) interval or (Q-A″max) as a function of (Q-A), the intervalbetween the Q-wave on EKG and the time of the femoral arterial pressurewave upstroke. A″max occurs at the maximum value of the secondderivative of pressure with respect to time. Both the arterial upstrokeand the instant of maximal upward pressure acceleration occur after theonset of systole. A″max always occurs after A since the pressure wavemust bottom out before it maximally accelerates. Inspection of FIG. 9shows that (Q-A″max) is greater than (Q-A) by a factor of 1.065 with anerror of 1%. The correlation coefficient between these two measures ofContractility is 0.965. The above correlation strongly indicates thatthe two intervals (Q-A) and (Q-A″max) are just different species of thegenus (E-M) and moreover that they are practically interchangeable.(Q-A) was described in the Lancet in 1974. (Q-A″max) has the advantagethat it is easily and accurately determinable when there is noise in thebaseline. Noisy baselines are part of the natural history of operatingrooms. Therefore, another aspect of the present invention provides amethod to determine the non-invasively measured cardiac parameters usingQ-A″max.

FIGS. 10 and 11 show the relation between the average rate of leftventricular outflow and the (E-M) interval. In both figures, the (E-M)interval used is the (Q-A″max) interval. The average rate of leftventricular (LV) outflow is just the Stroke Volume (SV) divided by theEjection Interval (EI) by definition. This yields a quantity, SV/EI,whose units are in cc/sec. FIGS. 10 and 11 respectively describe theabove relation for Pig 1 and Pig 2. It is clear that, in both cases, theaverage systolic ejection outflow rate is linearly proportional to ‘e’raised to the power of the reciprocal of the E-M interval. In Pig 1, thelinear correlation coefficient is 0.98. In Pig 2, the linear correlationcoefficient is 0.95. Therefore, it is another aspect of the presentinvention to provide a correlation between SV/EI and E-M interval, whichequals to Q-A″max.

In FIGS. 12 and 13, the previously described relation is solved for theStroke Volume (SV). If we can predict the average left ventricularoutflow rate, SV/EI on the basis of the (E-M) interval, then multiplyingboth sides of the equation by EI will give us a quantity which willtrack in a useful way with SV. FIG. 12 is a plot of Stroke Volumeagainst EI*exp(1/E-M) in Pig 1. FIG. 13 is the same plot of the datafrom Pig 2. In both plots, the (Q-A″max) interval was used for the (E-M)interval. Pig 1 data yields a linear correlation coefficient of 0.825.Pig 2 data yields a linear correlation coefficient of 0.944.Significantly, the product literature for Hemosonic esophageal Dopplerdevice indicates that its Cardiac Output measurements correlate withthermodilution measurements with a linear correlation coefficient of0.80. At least one investigator indicated the correlation at 0.90 (seeKlein, G., M.D., Emmerich, M., M.D., Clinical Evaluation of Non-invasiveMonitoring Aortic Blood Flow, (ABF) by a TransesophagealEcho-Doppler-Device. Anesthesiology 1998; V89 No. 3A: A953, op. cit.).The Hemosonic manual indicates that CO determined by the device isaccurate to +/−15%.

Correlations Between Parameters in One Dimension

In addition to the excellent correlation between the vector of theinvasively measured hemodynamic states and the vector of thenon-invasively measured hemodynamic states, the correlations betweeneach parameter of invasively measured hemodynamic states and itsrespective non-invasive counterpart are demonstrated separately in thefollowing discussions.

FIG. 14 shows the correlation of the average systolic outflow rate,SV/EI as a GV function of the natural log of the maximum value of thefirst derivative of left ventricular pressure with respect to time,ln(dP/dtmax) in Pig 1. This relation shows a linear correlationcoefficient, R of 0.9499 and a probability that the relation is due torandom chance, (P value) of <0.0001. The same relation is shown for Pig2 in FIG. 15. The value of R is 0.88304, and P<0.0001. Since SV/EIcorrelates linearly with exp(1/(E-M)) as shown in FIGS. 10 and 11, exp(1/(E-M)) must correlate linearly with respect to ln (dP/dtmax) as well.

This linear correlation between exp (1/E-M)) and dP/dtmax is furtherexperimentally demonstrated respectively in FIGS. 16 and 17 for Pigs 1and 2. FIG. 16 illustrates linear relation between 1/(Q-A″max) andln(dP/dtmax) in the Pig 1 experiments. The linear correlation as shownin FIG. 16 has a correlation coefficient of R=0.97472, and a P value of<0.0001. For Pig 2, R is 0.96009, and P is less than 0.0001 as shown inFIG. 17. These data show, for the first time, the exponential relationbetween dP/dtmax and 1/(Q-A″max) unlike the relationship betweendP/dtmax and Q-A implied by Jackson. (see Jackson, D. M., M.D, A SimpleNon-Invasive Technique for Measuring Cardiac Contractility, [Letter].Lancet 1974; ii:1457). This exponential relationship between dP/dtmaxand 1/(Q-A″max) provides a correlation that is necessary for the methodof the present invention of converting the non-invasively measuredparameters into the invasively measured parameters that are responsiveto medicines.

Jackson spoke of “changes in the Q-A interval” and not in terms of thereciprocal of the Q-A interval, 1/(Q-A). His data are plotted in termsof Δ(Q-A) or the difference in (Q-A) compared to the baseline. 1/(Q-A)is a more physiologically meaningful quantity.

The distance L from the aortic valve to the arterial pressure transducercatheter in the femoral artery stays constant throughout the experiment.Then, by definition, the quantity L/(Q-A) is the velocity ofelectromechanical transduction and propagation of the pulse wave downthe arterial tree. We can call L/(Q-A) the Velocity of transduction andpropagation or Vtp. Since L is a constant, 1/(Q-A) varies linearly withVtp. In other words,

Vtp=L[1/(Q-A)]  Eq. 24

So, 1/(Q-A) and by inference (1/(Q-A″max) give us a linear handle on thevelocity of electromechanical transduction and elastic wave propagation.

FIGS. 19, 20, and 21 describe certain experimental data from Pig 2. FIG.18 shows the relation between LVEDP and (EI*MAP*(Q-A″max)) for Pig 1.The linear correlation coefficient R=0.92088, and P<0.0001 for FIG. 18.FIG. 19 describes certain data in the Pig 2 experiments, but only showseleven of fifteen points. The points in FIG. 19 represent experiments 5through 15, which were performed serially in time. For these 11time-contiguous data points, R=0.94043, and P<0.0001. FIG. 20 shows datafor the experiments 1 through 4 of Pig 2. These first four experimentsyield R at 0.98471 and P at 0.01529. However, when all fifteen of thedata points from Pig 2 are shown together as in FIG. 21, the correlationcoefficient falls to 0.7067 with P at 0.0032.

FIGS. 22 and 23 respectively describe experimental data for Pig 1 andPig 2. FIG. 22 presents the double product [MAP*(Q-A″max)] as a functionof the Systemic Vascular Resistance, SVR. In the Pig 1 data, thecorrelation R=0.81052, and P<0.0001. In the Pig 2 data, R=0.93254, andP<0.0001.

In a second embodiment of the present invention, LVEDP, SVR and dP/dtmaxare correlated with the non-invasively measured results according toEqs. 19, 20 and 21. The correlations have led to the followingprocessing results and conclusions. There are close correlations betweenthe invasively measured parameters and the parameters derived fromnon-invasive measurement through another correlation method according tothe present invention. An improved formulation is disclosed for anon-invasively derived analogue of Left Ventricular End-DiastolicPressure (LVEDP). In diastole, the left ventricular pressure is anexponential function of left ventricular volume, and this relation holdsat any point during the diastolic filling interval includingend-diastole. Therefore, LVEDP is an exponential function of LeftVentricular End-Diastole Volume (LVEDV).

Let DI be the diastolic filling interval. To a reasonable approximation,

DI=T−EI  Eq. 25

where T is the time period of the cardiac cycle. T is non-invasivelyobtained by measuring the time interval between R-waves in the EKG andis linearly proportional to the reciprocal of the heart rate (HR) inbeats per minute. That is,

T=(1/HR)*60 sec/min  Eq. 26

This approximation ignores the time required for isovolumic contractionand relaxation. Since these intervals are relatively small fractions ofany cardiac cycle, the approximation is a useful one for the presentinvention. In addition to the method of obtaining DI as shown in Eq. 26,DI may be obtained by a 1 MHz Doppler ultrasound device placed on thesurface of the patient's chest just over the left ventricle. Diastolicfilling has a characteristic low velocity blood flow that causes ananalogously low Doppler frequency shift. The duration of thischaracteristic low frequency Doppler shift substantially serves as anaccurate measure of DI. DI starts when the mitral valve opens, and DIends when the mitral valve slams shut. The Doppler device is relativelyexpensive and has the advantage for use with obese patients. In somesituations, DI is also optionally obtained with an ordinary stethoscopeor phonocardiogram. An ordinary stethoscope or phonocardiogram indicatesthat DI ends as it is marked by the first heart sound of the ‘lub’ inthe two sounds of the lub-dub'. In patients with certain pathology, an‘opening snap’ of the mitral valve is audible in the stethoscope.Perhaps a phonocardiogram shows when the mitral valve opens in mostpatients. Alternatively, the above mentioned fiberoptic sensor that isplaced upon the precordium of the chest obtains DI by measuring theduration of the low frequency vibrations in the amplitude of thefiberoptic light signal which are due to diastolic filling. Preferably,DI used in the present invention is derived according to Eq. 25

The relation between LVEDP and the quantity DI=T−EI is shown for Pig 1in FIG. 24 and for Pig 2 in FIG. 25. The relation is sigmoid, and isgiven by the Boltzman equation

(T−EI)={(A1−A2)/[1+exp({LVEDP−x0)}/dx)]}+A2  Eq. 27

where A1 and A2 are asymptotes, A2>A1. The quantity (T−EI) at x=x0 isthe average value of the two asymptotes and ‘dx’ is a proportionalityconstant which decreases with the steepness of the sigmoid rise.

The data in FIGS. 24 and 25 represents a full range of Contractility andSVR. The data in FIGS. 24 and 25 shows excellent agreement with Eq. 27.By fitting these data with Eq. 27, we obtain a Chi-Square of 0.0002 forthe fitting of the Pig 1 data and a Chi-Square of 0.00078 for the Pig 2data. The remarkably good agreement between Eq. 27 and the experimentaldata shows a relationship between LVEDP and exactly one non-invasivelymeasurable time interval. When EI or T by itself is plotted as afunction of LVEDP, a similar sigmoid distribution results. But theChi-Square values for the difference DI=T−EI as a function of LVEDP arelower than for either quantity alone. Ordinarily, when two quantitiesare subtracted, their errors would add. In this case, the error actuallydecreases when (T−EI) is used in preference to either EI or T alone. Thesigmoid relation between DI and LVEDP represents a completely newcorrelation between two parameters in mammalian hemodynamics.

Preferably, the above sigmoid relation is used to achieve a betteraccuracy when LVEDP is near x0 and T−EI is not near one of itsasymptote. More preferably, the predictively useful range of (T−EI) forpredicting LVEDP is where LVEDP lies between 7-11 mmHg in both pigexperiments. Interestingly, the inflection point in the sigmoid curve x0is 9.40+/−0.127 for Pig 1, and x0 is 9.14+/−0.286 for Pig 2. Thesevalues of x0 for the two pigs are equal within the limits ofexperimental error. It is anticipated that with a large population ofpigs, that the quantity x0 would lie in a steep bell curve distribution.As known in the relevant prior art, the above data from the pigexperiments is a solid foundation for having the substantially similarcorrelation between the non-invasively measured data and the invasivelymeasured data from humans.

In order for the sigmoid relation shown in FIGS. 24 and 25 to be moreuseful for predicting LVEDP, T-EI is preferably made linear or at leastmonotonically increasing with respect to LVEDP and contains noasymptotes. This is accomplished by multiplying (T−EI) by the productMAP*(Q-A″max), where MAP is Mean Arterial Pressure and (Q-A″max) is thetime interval between the Q-wave on EKG and the time point of maximumvalue of the second derivative of arterial pressure with respect totime. In this example, the arterial pressure is measured in the femoralartery.

FIGS. 26 and 27 respectively show the quantity (T−EI)*MAP*(Q-A″max) as afunction of LVEDP in Pig 1 and Pig 2. In Pig 1, the linear correlationcoefficient R is 0.92586. In Pig 2, the linear correlation coefficient Requals to 0.8711. This represents a great improvement over the linearcorrelation coefficient of EI*MAP*(Q-A″max) vs. LVEDP, where the linearcorrelation coefficient R equals to 0.70677. Therefore,(T−EI)*MAP*(Q-A″max) is an improved non-invasively measured correlate ofLVEDP over EI*MAP*(Q-A″max).

The relations shown in FIGS. 26 and 27 are not perfectly linear. FIG. 26in particular has a modest exponential component. If we removed theexponential character from the Pig 1 results as shown in FIG. 26 byplotting (T−EI)*MAP*(Q-A″max) against an exponential function of LVEDP,we would get an even better linear correlation R=0.95702 as shown inFIG. 28. Similarly, if we removed the exponential character from the Pig2 results as shown in FIG. 27 by plotting (T−EI)*MAP*(Q-A″max) againstexp(LVEDP), we would get an improved linear correlation of R=0.91297 asshown in FIG. 29. It is reasonable that more and better data will leadto better correlations.

In addition to the correlations described above, additional correlationhas been developed between the non-invasively measured cardiacparameters and other cardiac parameters that are normally measuredinvasively or difficult to obtain.

1) Correlation Between SVRc and the Second Plurality of Non-InvasiveCardiac Parameters.

An alternative approximation to the Afterload is SVRc. SVRc is theSystemic Vascular Resistance defined only over the ejection interval(EI). That is precisely the interval over which resistance to flow isoffered by the resistance vessels at the arteriolar level of thecirculation. By contrast, SVR is defined over the entire cardiac cycle.Let SV be the stroke volume in cc over the ejection interval EI inseconds. Then by Ohm's Law,

SVRc=MAP/[SV/EI]  Eq. 29

We have already shown that the quantity [SV/EI] has a very high linearcorrelation with exp(1/Q-A″max), with R=0.97997 for Pig 1 and R=0.95425for Pig 2. Substituting exp(1/Q-A″max) for [SV/EI] in Eq. 29 gives us

SVRc∝MAP/[exp(1/Q-A″max)]  Eq. 30

where ‘∝’ indicates a linearly proportional relationship.

There is a practical problem with using Eq. 30 as an index of Afterload.As the denominator becomes small relative to the numerator, the randomerror in the quotient MAP/exp(1/(Q-A″max) becomes magnified in such away as to preclude deriving a continuous function at the right handportion of the curve where (Q-A″max) becomes large (that is, as(Q-A″max) becomes large and exp(1/Q-A″max) tends toward 1). This problemis easily fixed simply by adding a constant K in the denominator. K mustbe, sufficiently large that the denominator ‘[K+exp(1/Q-A″max)]’ is onthe same order of magnitude as the numerator, MAP over the physiologicrange. We can write

SVRc∝MAP/[K+exp(1/Q-A″max)]  Eq. 31

SVRc=A1*MAP/[K+exp(1/Q-A″max)]+A2  Eq. 31a

Where K, A1 and A2 are empirical proportionality constants. The units ofK are sec⁻¹. For Pig 1, K=400 sec⁻¹, and for Pig 2, K=70 sec⁻¹. Eq. 31shows that SVRc can be derived from non-invasively measured results MAP,Q-A″max.

FIGS. 30 and 31 plot the invasively measured SVRc againstMAP/[K+exp(1/Q-A″max)], which is essentially the SVRc derivednon-invasively respectively for Pig 1 and Pig 2. The correlationcoefficient R equals to 0.959 and 0.9648 respectively for Pig 1 and Pig2. The above correlations strongly confirm that SVRc is indeed derivedfrom MAP and Q-A″max, both of which are measured without invasion of thepatient.

To further confirm the practicality and effectiveness of the correlationin the second preferred embodiment according to the present invention,the vector trajectories in three dimensional hemodynamic invasive vectorspace are compared with vector trajectories in the Non-Invasive Space.The invasive spaces are shown in FIGS. 32, 34, and 36. Non-invasivespaces are shown in FIGS. 33, 35, and 37. Each of pairs of FIGS. 32 and33, 34 and 35; 36 and 37 describes the same events in the same pig witha vector obtained either invasively or non-invasively. Therefore, inanother embodiment, the plurality of invasive cardiac analogues arerepresented by LVEDP, SVRc and dP/dtmax, which are respectiveapproximation to P, A and C. Among the plurality of invasive cardiacanalogues, SVRc is defined by Eq. 31a.

FIGS. 32 and 33 represent experimental data for Pig 1. In FIG. 32, theinvasive Afterload is represented on the ‘y’ axis as SVRc=MAP/[SV/EI],that is the mean arterial pressure divided by the average systolicejection rate. In FIG. 33, the non-invasive Afterload is represented onthe ‘y’ axis by SVRc α MAP/[K+exp(1/Q-A″max)]. Also in FIG. 33, thenon-invasive Preload is represented as [(T−EI)*MAP*(Q-A″max)]. Thehomology between FIGS. 32 and 33 is striking and clearly visualized by askilled person in the art.

FIGS. 34 and 35 represent experimental data for Pig 2. In FIG. 34, theinvasive Afterload is represented on the ‘y’ axis as SVRc=MAP/[SV/EI],that is the mean arterial pressure divided by the average systolicejection rate. In FIG. 35, the non-invasive Afterload is represented onthe ‘y’ axis by SVRc α MAP/[K+exp(1/Q-A″max)]. Also in FIG. 35, thenon-invasive Preload is represented as [(T−EI)*MAP*(Q-A″max)]. Thehomology between FIGS. 34 and 35 is striking and clearly visualized by askilled person in the art.

FIGS. 36 and 37 also represent experimental data for Pig 2. In FIG. 36,the invasive Afterload is represented on the ‘y’ axis as SVR obtainedfrom the Swan-Ganz data. In FIG. 37, the non-invasive Afterload isrepresented on the ‘y’ axis by MAP*(Q-A″max), which is the non-invasiveanalogue of SVR. The non-invasive Preload is represented as[(T−EI)*MAP*(Q-A″max)]. The homology between FIGS. 36 and 37 is strikingand is visualized by a skilled person in the art.

2) An Improved Index of Left Ventricular Ischemia

The average compliance CP of the left ventricle over the diastolicfilling interval is

CP=ΔV/ΔP  Eq. 32

where ΔV is just the stroke volume SV. In the steady state, the volumeof blood filling the left ventricle equals the volume of blood ejectedfrom it. But since SV α EI*(exp(1/Q-A″max)) as shown previously, wesubstitute SV for ΔV in the numerator of Eq. 32. AP in diastole is theLeft Ventricular End-Diastolic Pressure (LVEDP) minus the LV Pressure atthe end of isovolumic relaxation (Peivr). Thus, ΔP=LVEDP−Peivr. SincePeivr is always going to be a very low number or close to zero, forpractical purposes, we can ignore it. Assuming that Peivr is near zero,and substituting EI*exp(1/Q-Amax) for ΔV and (T−EI)*MAP*(Q-A″max) for ΔPin Eq. 32, we get

CP∝[EI*exp(1/Q-A″max)]/[(T−EI)*MAP*(Q-A″max)]  Eq. 33

Rearranging terms, we get

CP∝[EI/(T−EI)]*[exp(1/Q-A″max)/(MAP*Q-A″max)]  Eq. 34

or CP=A3*[EI/(T−EI)]*[exp(1/Q-A″max)/(MAP*Q-A″max)]+A4  Eq. 34a

where A3 and A4 are empirical proportionality constants.

In an ischemic event, there should be a sudden decrease in the value ofthe average compliance CP that precedes the onset of regional wallmotion abnormalities on 2-D echocardiography, which in turn will precedethe onset of S-T segment changes on EKG. The above diagnostic feature ofthe average compliance CP would make non-invasive left ventriculardiastolic compliance measurement far superior and more sensitive to anydiagnostic tool that is now available for ischemia detection andtreatment efficacy monitoring. Therefore, a preferred method ofpredicting an ischemic event according to the present invention detectsischemic events well ahead of other currently available monitors bymeasuring T, EI, MAP and Q-A″max. Therefore, in another aspect, thepresent invention provides a method, a system and a correlation toderive at least one clinically useful cardiac parameter or invasivecardiac analogue from a plurality of predetermined non-invasivelymeasured parameters such as T, EI, MAP, Q-D″(t) max, and Q-A″max. Theclinically useful cardiac parameters include CP, which is useful inpredicting an ischemic event; and P, A, C, analogues thereof andapproximations thereof, approximations thereof, which are useful inmonitoring a patient's cardiac state and in determining the need ofcardiac medicines such as dobutamine, nitroglycerine, phenylephrine,fluids, diuretics, pressors, afterload reducers, anesthetics, inotropesand negative inotropes.

Although the inventor does not want to limit the predictability of[EI/(T−EI)]*[exp(1/Q-A″max)/(MAP*Q-A″max)] according to the presentinvention to a particular theory, a possible explanation for detectingan ischemic event is as follows. An acute ischemic event typically doesnot manifest itself in exactly one cardiac function parameter. Rather,it is a rapidly emerging gestalt that affects all four parametersincluding heart rate, Preload, Afterload, and Contractility in acharacteristic way over a characteristic time course. During an ischemicevent, myocardial oxygen demand exceeds the supply. Now, there is notenough ATP being created by the oxygen requiring respiratory enzymes forthe oxidation of glucose substrate. The missing ATP is needed to breakthe cross-links that form between actin and myosin fibrils as a resultof muscle contraction. Muscle relaxation is an energy-requiring andoxygen-requiring process. Without sufficient oxygen, the actin-myosincross-links remain in place at the end of diastole. The remaining,unbroken cross-links cause the left ventricle muscle to stiffen duringfilling. In other words, left ventricular diastolic compliancedecreases.

Concomitant with the ischemic event is a decrease in Contractility. Atsame time, stroke volume and cardiac output fall. Afterload increases asthe organism responds to sympathetic nerve signals by increasing thesystemic vascular resistance (SVR) by selectively constricting smoothmuscular arterioles. The increase in SVR compensates in part for thefall in vital end-organ perfusion pressure caused by the fall in cardiacoutput. By globally increasing SVR by selective vasoconstriction ofblood flow to non-critical organ systems such as the gut, liver andmuscles, more of the falling and increasingly scarce global cardiacoutput can be shunted to the brain, heart, and lungs. In short order,Preload increases causing the Starling mechanism of the myocardium tocome into play, which compensates for the decrease in stroke volume byincreasing it. This increase in Preload occurs simply because, with adecrease in contractility, a smaller fraction of the Left VentricularEnd-diastolic Volume is ejected with each stroke, causing blood volumeto accumulate in the Left Ventricle with diastolic filling, therebyincreasing LVEDP. This distension of the left ventricle in diastolicfilling is precisely what we mean by an increase in Preload. With anischemic event, a choreographed increase in LVEDP, an increase in SVR,and a decrease in Contractility occur almost simultaneously.

Now, referring to Eq. 34, the right hand member contains the termexp(1/(Q-A″max)) in the numerator, which is proportional toContractility. Therefore, if the compliance CP suddenly decreases, Eq.34 tells us that Contractility also linearly decreases. The right handmember of Eq. 34 contains [T−EI)*MAP*(Q-A″max)], which is proportionalto Preload, in the denominator. Therefore, by Eq. 34, the compliance CPvaries inversely with Preload. If the compliance CP suddenly decreasesdue to ischemia, Preload inversely increases. The right hand member ofEq. 34 contains the expression [exp(1/Q-A″max)/MAP], which isproportional to (1/SVRc). SVRc then, by Eq. 34, varies inversely withdiastolic compliance CP. So if CP were to decrease suddenly due toischemia, SVRc would increase. But these changes in LVEDP, dP/dtmax, andSVRc that are described by Eq. 34 in the event of a sudden decrease inLV compliance are precisely the choreographed changes in Preload,Afterload, and Contractility that occur in the event of ischemia. Insummary, if the compliance CP of the Left Ventricle declines suddenlydue to ischemia, this entails an increase in LVEDP, an increase in SVRc,and a sudden decrease in Contractility, which is consistent with ourphysiological knowledge. What is significant here is that thisparticular piece of physiologic knowledge can be shown to be a purelylogical consequence of the definition of average diastolic compliance,CP, and Eq. 34, which is a relation between average diastolic complianceand a plurality of pre-determined non-invasive cardiac parameters. Thesenon-invasive parameters, in turn, determine the locus of a point in thenon-invasive hemodynamic vector space, ‘N’.

Another approach to the problem of finding a non-invasively measuredindex of cardiac ischemia, that is calculated and displayed inreal-time, is as follows. We define a vector function of three mutuallyorthogonal variables,

Xi={LVEDP,SVRc,1/(ln(dP/dtmax))}  Eq. 36

We then express the above vector in terms of its non-invasivecorrelates,

Xn={[(T−EI)*MAP*(Q-A″max],[MAP/(K+exp(1/Q-A″max))],(Q-A″max)}  Eq. 37

Then we find the norm of the above vector, that is, the square root ofthe dot product of the vector with itself. This will gives us anischemia function ‘I’, which will abruptly increase during an episode ofischemia, where

I={(T−EI)²*MAP²*(Q-A″max)²+MAP²/{K+exp(1/Q-A″max)}²+(Q-A″max)²}^(1/2)  Eq. 38

While there have been described what are believed to be the preferredembodiments of the present invention, those skilled in the art willrecognize that other and further changes and modifications may be madethereto without departing from the spirit of the invention, and it isintended to claim all such changes and modifications as fall within thetrue scope of the invention.

1. A method of monitoring cardiac parameters, comprising the steps of:non-invasively measuring a plurality of predetermined non-invasivecardiac parameters from a subject; and converting the non-invasivecardiac parameters into a plurality of invasive cardiac analogues basedupon a set of predetermined conversion equations.
 2. The method ofmonitoring cardiac parameters according to claim 1 wherein the subjectis a human.
 3. The method of monitoring cardiac parameters according toclaim 1 wherein the subject is an animal.
 4. The method of monitoringcardiac parameters according to claim 1 wherein the predeterminednon-invasive cardiac parameters include heart rate as denoted by HR,ejection interval as denoted by EI, mean arterial pressure as denoted byMAP and electrical-mechanical interval as denoted by E-M, which is aninterval between an electrical event E and a mechanical event M.
 5. Themethod of monitoring cardiac parameters according to claim 4 wherein thepredetermined invasive cardiac analogues include preload as denoted byP, afterload as denoted by A and contractility as denoted by C.
 6. Themethod of monitoring cardiac parameters according to claim 5 wherein thepredetermined conversion equations includeP=k1(EI*MAP*E-M)+c1,A=k2(MAP*E-M)+c2, andln(C)=k3(1/E-M)+c3 where k1, k2, k3, c1, c2 and c3 are empiricalproportionality constants.
 7. The method of monitoring cardiacparameters according to claim 6 wherein the M in the E-M is defined as atime when a second derivative with respect to time, M″(t), reaches amaximum value.
 8. The method of monitoring cardiac parameters accordingto claim 6 wherein the electrical event in determining the E-M isselected from the group consisting of a Q-wave, a R-wave, an S-wave, andan artificial ventricular pacemaker spike.
 9. The method of monitoringcardiac parameters according to claim 6 wherein the electrical event indetermining the E-M interval is determined by double differentiating anEKG voltage curve which corresponds to ventricular depolarization, V(t),with respect to time and defining the electrical event as a time whenV″(t) reaches a maximum positive value.
 10. The method of monitoringcardiac parameters according to claim 6 wherein the second mechanicalevent in determining the E-M is selected from the group consisting ofarterial blood pressure and flow velocity upstroke.
 11. The method ofmonitoring cardiac parameters according to claim 5 wherein thepredetermined conversion equations includeP=k1′((T−EI)*MAP*E-M)+c1′,A=k2′(MAP*E-M)+c2′, andln(C)=k3′(1/E-M)+c3′ where k1′, k2′, k3′, c1′, c2′ and c3′ are empiricalproportionality constants for a particular one of the subjects and T isa time period of the cardiac cycle for the particular one of thesubjects.
 12. The method of monitoring cardiac parameters according toclaim 11 wherein the M in the E-M is defined as a time when a secondderivative with respect to time, M″(t), reaches a maximum value.
 13. Themethod of monitoring cardiac parameters according to claim 11 whereinthe electrical event in determining the E-M is selected from the groupconsisting of a Q-wave, a R-wave, an S-wave, and an artificialventricular pacemaker spike.
 14. The method of monitoring cardiacparameters according to claim 11 wherein the electrical event indetermining the E-M interval is determined by double differentiating anEKG voltage curve which corresponds to ventricular depolarization, V(t),with respect to time and defining the electrical event as a time whenV″(t) reaches a maximum positive value.
 15. The method of monitoringcardiac parameters according to claim 13 wherein a mechanical event indetermining the E-M includes a time of an arterial blood pressureupstroke as denoted by TA and a time of a blood flow velocity upstrokeas denoted by TF.
 16. The method of monitoring cardiac parametersaccording to claim 5 wherein the predetermined conversion equationsincludeP=k1((T−EI)*MAP*E-M)+c1′,A=k2′*MAP/[K+exp(1/E-M)]+c2′, andln(C)=k3′(1/E-M)+c3′ where k1′, k2′, k3′, c1′, c2′, c3′, and K areempirical proportionality constants for a particular one of the subjectsand T is a time period of the cardiac cycle for the particular one ofthe subjects.
 17. The method of monitoring cardiac parameters accordingto claim 4 wherein the EI is measured by placing a Doppler ultrasounddevice over the suprasternal notch near the ascending aorta.
 18. Themethod of monitoring cardiac parameters according to claim 4 wherein theelectrical event E in the E-M is determined by electrocardiograph asdenoted by EKG.
 19. The method of monitoring cardiac parametersaccording to claim 4 wherein the E-M is determined by electrocardiographas denoted by EKG and by placing a Doppler ultrasound device over amajor artery.
 20. The method of monitoring cardiac parameters accordingto claim 4 wherein the E-M is determined by electrocardiograph asdenoted by EKG and by placing a fiberoptic device over a major artery.21. The method of monitoring cardiac parameters according to claim 1further comprising the step of displaying the invasive cardiac analoguesin three dimensional coordinate space that is defined by a first axisindicative of the P, a second axis indicative of the A and a third axisindicative of the C.
 22. The method of monitoring cardiac parametersaccording to claim 21 further comprising an additional step ofdisplaying a three dimensional object defining a safe zone indicative ofa safe hemodynamic state.
 23. The method of monitoring cardiacparameters according to claim 22 wherein the first axis, the secondaxis, the third axis and the three dimensional object are each displayedwith a predetermined color.
 24. The method of monitoring cardiacparameters according to claim 1 further comprising an additional step ofdisplaying a vector cross product indicating an amount of physiologicstress.
 25. The method of monitoring cardiac parameters according toclaim 1 further comprising an additional step of determining a fitnesslevel of the subject based upon the invasive cardiac analogues.
 26. Themethod of monitoring cardiac parameters according to claim 1 furthercomprising the step of determining management of an anesthetic-relatedprocedure of the subject based upon the invasive cardiac analogues. 27.The method of monitoring cardiac parameters according to claim 1 furthercomprising the step of determining an abnormal cardiac condition of thesubject based upon the invasive cardiac analogues.
 28. The method ofmonitoring cardiac parameters according to claim 1 further comprisingthe step of transferring the non-invasive cardiac parameters from onelocation to another location before converting the non-invasive cardiacparameters into the invasive cardiac analogues.
 29. The method ofmonitoring cardiac parameters according to claim 28 further comprisingthe step of evaluating a cardiac condition of the subject based upon theinvasive cardiac analogues.
 30. A system for monitoring cardiacparameters comprising: a non-invasive cardiac parameter measuring unitfor non-invasively measuring a plurality of predetermined non-invasivecardiac parameters from a subject; and a conversion unit connected tosaid non-invasive cardiac parameter measuring unit for converting thenon-invasive cardiac parameters into a plurality of invasive cardiacanalogues based upon a set of predetermined conversion equations. 31.The system for monitoring cardiac parameters according to claim 30wherein said non-invasive cardiac parameter measuring unit measures thepredetermined non-invasive cardiac parameters from a human.
 32. Thesystem for monitoring cardiac parameters according to claim 30 whereinsaid non-invasive cardiac parameter measuring unit measures thepredetermined non-invasive cardiac parameters from an animal.
 33. Thesystem for monitoring cardiac parameters according to claim 30 whereinsaid non-invasive cardiac parameter measuring unit further comprises aheart rate monitor for measuring heart rate as denoted by HR, avibration sensing device for measuring an ejection interval as denotedby EI and a mechanical event M of an electrical-mechanical interval asdenoted by E-M, a blood pressure measuring device for measuring meanarterial pressure as denoted by MAP and an electrocardiogram measuringdevice for measuring an electrical event E of the electrical-mechanicalinterval.
 34. The system for monitoring cardiac parameters according toclaim 33 wherein said vibration sensing device comprises at least one ofa Doppler ultrasound device and a fiber optic device.
 35. The system formonitoring cardiac parameters according to claim 30 wherein saidconversion unit outputs the predetermined invasive cardiac analoguesincluding preload as denoted by P, afterload as denoted by A andcontractility as denoted by C.
 36. The system for monitoring cardiacparameters according to claim 35 wherein said conversion unit determinesthe P, the A and the C based upon the predetermined conversion equationsincludingP=k1(EI*MAP*E-M)+c1,A=k2(MAP*E-M)+c2, andln(C)=k3(1/E-M)+c3 where k1, k2, k3, c1, c2 and c3 are empiricalproportionality constants.
 37. The system for monitoring cardiacparameters according to claim 35 wherein said conversion unit determinesthe P, the A and the C based upon the predetermined conversion equationsincludingP=k1′((T−EI)*MAP*E-M)+c1′,A=k2′(MAP*E-M)+c2′, andln(C)=k3′(1/E-M)+c3′ where k1′, k2′, k3′, c1′, c2′ and c3′ are empiricalproportionality constants for a particular one of the subjects and T isa time period of the cardiac cycle for the particular one of thesubjects.
 38. The system for monitoring cardiac parameters according toclaim 35 wherein said conversion unit determines the P, the A and the Cbased upon the predetermined conversion equations includingP=k1′(DI*MAP*E-M)+c1′,A=k2′(MAP*E-M)+c2′, andln(C)=k3′(1/E-M)+c3′ where k1′, k2′, k3′, c1′, c2′ and c3′ are empiricalproportionality constants for a particular one of the subjects, DI is adiastolic filling interval, and T is a time period of the cardiac cyclefor the particular one of the subjects.
 39. The system for monitoringcardiac parameters according to claim 37 wherein said conversion unitobtains the mechanical event M in the E-M by determining a time when asecond derivative with respect to time, M″ (t) reaches a maximum value.40. The system for monitoring cardiac parameters according to claim 39wherein said non-invasive cardiac parameter measuring unit measures anelectrical event in determining the E-M, the electrical event beingselected from the group consisting of a Q-wave as denoted by Q, a R-waveas denoted by R, a S-wave as denoted by S and an artificial ventricularpacemaker spike.
 41. The system for monitoring cardiac parametersaccording to claim 39 wherein said non-invasive cardiac parametermeasuring unit measures an electrical event in determining the E-Minterval by double differentiating an EKG voltage curve whichcorresponds to ventricular depolarization, V(t), with respect to timeand defining the electrical event as a time when V″(t) reaches a maximumpositive value.
 42. The system for monitoring cardiac parametersaccording to claim 40 wherein said non-invasive cardiac parametermeasuring unit measures the mechanical event in determining the E-M, thesecond mechanical event including a time of an arterial blood pressureupstroke as denoted by TA and a time of a flow velocity upstroke asdenoted by TF.
 43. The system for monitoring cardiac parametersaccording to claim 35 wherein said conversion unit determines the P, theA and the C based upon the predetermined conversion equations includingP=k1′((T−EI)*MAP*E-M)+c1′,A=k2′*MAP/[K+exp(1/E-M)]+c2′, andln(C)=k3′(1/E-M)+c3′ where k1′, k2′, k3′, c1′, c2′, c3′, and K areempirical proportionality constants for a particular one of the subjectsand T is a time period of the cardiac cycle for the particular one ofthe subjects.
 44. The system for monitoring cardiac parameters accordingto claim 30 further comprising a display unit connected to saidconversion unit for displaying the invasive cardiac analogues in threedimensional coordinate space that is defined by a first axis indicativeof the P, a second axis indicative of the A and a third axis indicativeof the C.
 45. The system for monitoring cardiac parameters according toclaim 44 wherein said display unit additionally displays a threedimensional object defining a safe zone indicative of a safe hemodynamicstate.
 46. The system for monitoring cardiac parameters according toclaim 45 wherein said display unit displays the first axis, the secondaxis, the third axis and the safe zone respectively in a predeterminedcolor.
 47. The system for monitoring cardiac parameters according toclaim 45 wherein said display unit additionally displays a vector crossproduct indicative of an amount of physiologic stress.
 48. The systemfor monitoring cardiac parameters according to claim 30 furthercomprising a determination unit connected to said conversion unit fordetermining a fitness level of the subject based upon the invasivecardiac analogues.
 49. The system for monitoring cardiac parametersaccording to claim 30 further comprising a determination unit connectedto said conversion unit for determining management of ananesthetic-related procedure of the subject based upon the invasivecardiac analogues.
 50. The system for monitoring cardiac parametersaccording to claim 30 further comprising a determination unit connectedto said conversion unit for determining an abnormal cardiac condition ofthe subject based upon the invasive cardiac analogues.
 51. The systemfor monitoring cardiac parameters according to claim 30 furthercomprising a data communication unit connected to said non-invasivecardiac parameter measuring unit at one location for transferring thenon-invasive cardiac parameters to said conversion unit at anotherlocation before converting the non-invasive cardiac parameters into theinvasive cardiac analogues.
 52. The system for monitoring cardiacparameters according to claim 51 wherein said data communication unittransfers the non-invasive cardiac parameters to said conversion unitvia the Internet.
 53. The system for monitoring cardiac parametersaccording to claim 51 wherein said data communication unit transfers thenon-invasive cardiac parameters to said conversion unit viatelecommunication.
 54. The system for monitoring cardiac parametersaccording to claim 30 wherein said non-invasive cardiac parametermeasuring unit is portable.
 55. The system for monitoring cardiacparameters according to claim 30 wherein said conversion unit isretrofitted to an existing one of said non-invasive cardiac parametermeasuring unit.
 56. A system for retrofitting existing non-invasivecardiac parameter measuring devices to generate invasive cardiacanalogues, comprising: an interface unit for receiving predeterminednon-invasive cardiac parameters of a subject from the existingnon-invasive cardio monitoring devices; and a conversion unit connectedto said interface unit for converting the non-invasive cardiacparameters into a plurality of the invasive cardiac analogues based upona set of predetermined conversion equations.
 57. The system forretrofitting existing non-invasive cardio monitoring devices accordingto claim 56 wherein the existing non-invasive cardiac parametermeasuring devices measure the predetermined non-invasive cardiacparameters from a human.
 58. The system for retrofitting existingnon-invasive cardio monitoring devices according to claim 56 wherein theexisting non-invasive cardiac parameter measuring devices measure thepredetermined non-invasive cardiac parameters from an animal.
 59. Thesystem for retrofitting existing non-invasive cardio monitoring devicesaccording to claim 56 wherein the existing non-invasive cardiacparameter measuring unit further comprises a heart rate monitor formeasuring heart rate as denoted by HR, a vibration sensing device formeasuring an ejection interval as denoted by EI and a mechanical event Mof an electrical-mechanical interval as denoted by E-M, a blood pressuremeasuring device for measuring mean arterial pressure as denoted by MAPand an electrocardiogram measuring device for measuring an electricalevent E of the electrical-mechanical interval.
 60. The system forretrofitting existing non-invasive cardio monitoring devices accordingto claim 56 wherein the existing non-invasive cardiac parametermeasuring unit further comprises a heart rate monitor for measuringheart rate as denoted by HR, a vibration sensing device for measuring anejection interval as denoted by EI, a plethysmographic device formeasuring a mechanical event M of an electrical-mechanical interval asdenoted by E-M, a blood pressure measuring device for measuring meanarterial pressure as denoted by MAP and an electrocardiogram measuringdevice for measuring an electrical event E of the electrical-mechanicalinterval.
 61. The system for retrofitting existing non-invasive cardiomonitoring devices according to claim 59 wherein said vibration sensingdevice comprises at least one of a Doppler ultrasound device and a fiberoptic device.
 62. The system for retrofitting existing non-invasivecardio monitoring devices according to claim 59 wherein said conversionunit outputs the predetermined invasive cardiac analogues includingpreload as denoted by P, afterload as denoted by A and contractility asdenoted by C.
 63. The system for retrofitting existing non-invasivecardio monitoring devices according to claim 62 wherein said conversionunit determines the P, the A and the C based upon the predeterminedconversion equations includingP=k1(EI*MAP*E-M)+c1,A=k2(MAP*E-M)+c2, andln(C)=k3(1/E-M)+c3 where k1, k2, k3, c1, c2 and c3 are empiricalproportionality constants.
 64. The system for retrofitting existingnon-invasive cardio monitoring devices according to claim 62 whereinsaid conversion unit determines the P, the A and the C based upon thepredetermined conversion equations includingP=k1′((T−EI)*MAP*E-M)+c1′,A=k2′(MAP*E-M)+c2′, andln(C)=k3′(1/E-M)+c3′ where k1′, k2′, k3′, c1′, c2′ and c3′ are empiricalproportionality constants for a particular one of the subjects and T isa time period of the cardiac cycle for the particular one of thesubjects.
 65. The system for retrofitting existing non-invasive cardiomonitoring devices according to claim 62 wherein said conversion unitdetermines the P, the A and the C based upon the predeterminedconversion equations includingP=k1′(DPMAP*E-M)+c1′,A=k2′(MAP*E-M)+c2′, andln(C)=k3′(1/E-M)+c3′ where k1′, k2′, k3′, c1′, c2′ and c3′ are empiricalproportionality constants for a particular one of the subjects and T isa time period of the cardiac cycle for the particular one of thesubjects.
 66. The system for retrofitting existing non-invasive cardiomonitoring devices according to claim 64 wherein the M in the E-M isdefined as a time when a second derivative with respect to time, M″(t),reaches a maximum value.
 67. The system for retrofitting existingnon-invasive cardio monitoring devices according to claim 66 whereinsaid non-invasive cardiac parameter measuring unit measures anelectrical event in determining the E-M, the electrical event beingselected from the group consisting of a Q-wave as denoted by Q, a R-waveas denoted by R, a S-wave as denoted by S and an artificial ventricularpacemaker spike.
 68. The system for retrofitting existing non-invasivecardio monitoring devices according to claim 66 wherein saidnon-invasive cardiac parameter measuring unit measures an electricalevent in determining the E-M interval by double differentiating an EKGvoltage curve which corresponds to ventricular depolarization, V(t),with respect to time and defining the electrical event as a time whenV″(t) reaches a maximum positive value.
 69. The system for retrofittingexisting non-invasive cardio monitoring devices according to claim 66wherein said non-invasive cardiac parameter measuring unit measures themechanical event in determining the E-M, the mechanical event selectedfrom the group consisting of a time of an arterial blood pressureupstroke as denoted by TA and a time of a flow velocity upstroke asdenoted by TF.
 70. The system for retrofitting existing non-invasivecardio monitoring devices according to claim 62 wherein said conversionunit determines the P, the A and the C based upon the predeterminedconversion equations includingP=k1′((T−EI)*MAP*E-M)+c1′,A=k2′*MAP/[K+exp(1/E-M)]+c2′, andln(C)=k3′(1/E-M)+c3′ where k1′, k2′, k3′, c1′, c2′, c3′, and K areempirical proportionality constants for a particular one of the subjectsand T is a time period of the cardiac cycle for the particular one ofthe subjects.
 71. The system for retrofitting existing non-invasivecardio monitoring devices according to claim 56 further comprising adisplay unit connected to said conversion unit for displaying theinvasive cardiac analogues in three dimensional coordinate space that isdefined by a first axis indicative of the P, a second axis indicative ofthe A and a third axis indicative of the C.
 72. The system forretrofitting existing non-invasive cardio monitoring devices accordingto claim 71 wherein said display unit additionally displays a threedimensional object defining a safe zone indicative of a safe hemodynamicstate.
 73. The system for retrofitting existing non-invasive cardiomonitoring devices according to claim 72 wherein said display unitdisplays the first axis, the second axis, the third axis and the safezone respectively in a predetermined color.
 74. The system forretrofitting existing non-invasive cardio monitoring devices accordingto claim 72 wherein said display unit additionally displays a vectorcross product indicative an amount of physiologic stress.
 75. The systemfor retrofitting existing non-invasive cardio monitoring devicesaccording to claim 56 further comprising a determination unit connectedto said conversion unit for determining a fitness level of the subjectbased upon the invasive cardiac analogues.
 76. The system forretrofitting existing non-invasive cardio monitoring devices accordingto claim 56 further comprising a determination unit connected to saidconversion unit for determining management of an anesthetic-relatedprocedure of the subject based upon the invasive cardiac analogues. 77.The system for retrofitting existing non-invasive cardio monitoringdevices according to claim 56 further comprising a determination unitconnected to said conversion unit for determining an abnormal cardiaccondition of the subject based upon the invasive cardiac analogues. 78.The system for retrofitting existing non-invasive cardio monitoringdevices according to claim 56 further comprising a data communicationunit connected to the existing invasive cardiac parameter measuringdevices at one location for transferring the non-invasive cardiacparameters to said interface unit at another location before convertingthe non-invasive cardiac parameters into the invasive cardiac analogues.79. The system retrofitting existing non-invasive cardio monitoringdevices according to claim 78 wherein said data communication unittransfers the non-invasive cardiac parameters to said interface unit viathe Internet.
 80. The system for retrofitting existing non-invasivecardio monitoring devices according to claim 78 wherein said datacommunication unit transfers the non-invasive cardiac parameters to saidinterface unit via telecommunication.
 81. The system for retrofittingexisting non-invasive cardio monitoring devices according to claim 56wherein the existing non-invasive cardio monitoring devices are allportable.
 82. A method of determining a patient's cardiac contractility,comprising the steps of: non-invasively measuring the patient'selectrocardiograph having a predetermined electrical wave; determining afirst point within a predetermined cardiac cycle based upon thepredetermined electrical wave; non-invasively measuring the patient'sarterial pressure with respect to time; determining a second point inthe predetermined cardiac cycle when a second derivative of apredetermined physiological function with respect to time reachesmaximum; and obtaining the cardiac contractility based upon the firstpoint and the second point.
 83. The method of determining a patient'scardiac contractility according to claim 82 wherein said predeterminedelectrical wave is selected from the group consisting of a Q-wave asdenoted by Q, a R-wave as denoted by R, a S-wave as denoted by S and anartificial ventricular pacemaker spike.
 84. The method of determining apatient's cardiac contractility according to claim 82 wherein theelectrical event in determining the E-M interval is determined by doubledifferentiating an EKG voltage curve which corresponds to ventriculardepolarization, V(t), with respect to time and defining the electricalevent as a time when V″(t) reaches a maximum positive value.
 85. Themethod of determining a patient's cardiac contractility according toclaim 82 wherein an improved cardiac contractility measure is furtherdefined as proportional to an exponential function of a reciprocal of atime interval defined between the first point and the second point. 86.A system for determining a patient's cardiac contractility, comprising:an electrocardiogram unit for non-invasively measuring the patient'selectrocardiograph having a predetermined electrical wave; an arterialpressure measuring unit for non-invasively measuring the patient'sarterial pressure with respect to time; and a determination unitconnected to said electrocardiogram unit and said arterial pressuremeasuring unit for determining a first point having a minimum within apredetermined cardiac cycle based upon the predetermined electrical waveand for determining a second point in the predetermined cardiac cyclewhen a second derivative of a predetermined physiological function withrespect to time reaches maximum based upon the patient's arterialpressure, said determination unit obtaining the cardiac contractilitybased upon the first point and the second point.
 87. The system fordetermining a patient's cardiac contractility according to claim 86wherein said predetermined electrical wave is selected from the groupconsisting of a Q-wave as denoted by Q, a R-wave as denoted by R, aS-wave as denoted by S and an artificial ventricular pacemaker spike.88. The system for determining a patient's cardiac contractilityaccording to claim 86 wherein the electrical event in determining theE-M interval is determined by double differentiating an EKG voltagecurve which corresponds to ventricular depolarization, V(t), withrespect to time and defining the electrical event as a time when V″(t)reaches a maximum positive value.
 89. The system for determining apatient's cardiac contractility according to claim 87 wherein animproved cardiac contractility measure is further defined asproportional to an exponential function of a reciprocal of a timeinterval defined between the first point and the second point.
 90. Amethod of monitoring an ischemic event, comprising the steps of:non-invasively measuring a plurality of predetermined non-invasivecardiac parameters from a subject; and converting the non-invasivecardiac parameters into a single invasive cardiac analogue indicative ofthe ischemic event based upon a predetermined conversion equation. 91.The method of monitoring an ischemic event according to claim 90 whereinthe subject is a human.
 92. The method of monitoring an ischemic eventaccording to claim 90 wherein the subject is an animal.
 93. The methodof monitoring an ischemic event according to claim 90 wherein thepredetermined non-invasive cardiac parameters include heart rate asdenoted by HR, ejection interval as denoted by EI, mean arterialpressure as denoted by MAP and electrical-mechanical interval as denotedby E-M, which is an interval between an electrical event E and amechanical event M.
 94. The method of monitoring an ischemic eventaccording to claim 93 wherein the predetermined conversion equationincludes CP=A3*[EI/(T−EI)]*[exp(1/E-M)/(MAP*E-M)]+A4, where T is cardiacperiod, which is obtained from the heartrate HR, A3 and A4 are empiricalproportionality constants and wherein the M in the E-M is furtherdefined as a time when, a second derivative of the M with respect totime, M″(t), reaches a maximum value.
 95. The method of monitoring anischemic event according to claim 94 wherein an electrical event indetermining the E-M is selected from the group consisting of a Q-wave asdenoted by Q, a R-wave as denoted by R, a S-wave as denoted by S and anartificial ventricular pacemaker spike.
 96. The method of monitoring anischemic event according to claim 94 wherein the electrical event indetermining the E-M interval is determined by double differentiating anEKG voltage curve which corresponds to ventricular depolarization, V(t),with respect to time and defining the electrical event as a time whenV″(t) reaches a maximum positive value.
 97. The method of monitoring anischemic event according to claim 94 wherein the mechanical eventselected from the group consisting of a time of an arterial bloodpressure upstroke as denoted by TA, a time of a flow velocity upstrokeas denoted by TF, and a time of plethysmographic upstroke as denoted byTOP.
 98. A system for monitoring an ischemic event, comprising: ameasuring unit for non-invasively measuring a plurality of predeterminednon-invasive cardiac parameters from a subject; and a converting unitconnected to said measuring unit for converting the non-invasive cardiacparameters into a single invasive cardiac analogue indicative of theischemic event based upon a predetermined conversion equation.
 99. Thesystem for monitoring an ischemic event according to claim 98 whereinthe subject is a human.
 100. The system for monitoring an ischemic eventaccording to claim 98 wherein the subject is an animal.
 101. The systemfor monitoring an ischemic event according to claim 98 wherein saidmeasuring unit measures the predetermined non-invasive cardiacparameters including heart rate as denoted by HR, ejection interval asdenoted by EI, mean arterial pressure as denoted by MAP andelectrical-mechanical interval as denoted by E-M.
 102. The system ofmonitoring an ischemic event according to claim 101 wherein thepredetermined conversion equation includesCP=A3*[EI(T−EI)]*[exp(1/E-M)/(MAP*E-M)]+A4, where T is cardiac period,which is obtained from the heartrate HR, A3 and A4 are empiricalproportionality constants and wherein the M in the E-M is furtherdefined as a time when, a second derivative of the M with respect totime, M″(t), reaches a maximum value.
 103. The system of monitoring anischemic event according to claim 102 wherein an electrical event indetermining the E-M is selected from the group consisting of a Q-wave asdenoted by Q, a R-wave as denoted by R, a S-wave as denoted by S and anartificial ventricular pacemaker spike.
 104. The system of monitoring anischemic event according to claim 102 wherein the electrical event indetermining the E-M interval is determined by double differentiating anEKG voltage curve which corresponds to ventricular depolarization, V(t),with respect to time and defining the electrical event as a time whenV″(t) reaches a maximum positive value.
 105. The system of monitoring anischemic event according to claim 103 wherein the second mechanicalevent selected from the group consisting of a time of an arterial bloodpressure upstroke as denoted by TA, a time of a flow velocity upstrokeas denoted by TF, and a time of plethysmographic upstroke as denoted byTOP.